Heart pump controller

ABSTRACT

A heart pump including first and second cavities, each cavity including a respective inlet and outlet, a connecting tube extending between the first and second cavities, an impeller including: a first set of vanes mounted on a first rotor in the first cavity portion; a second set of vanes mounted on a second rotor in the second cavity portion; and, a shaft connecting the first and second rotors, the shaft extending through the connecting tube, a drive for rotating the impeller and a magnetic bearing including at least one bearing coil for controlling an axial position of the impeller, at least one of the drive and magnetic bearing being mounted outwardly of the connecting tube, at least partially between the first and second cavity portions.

BACKGROUND OF THE INVENTION

The present invention relates to a controller for a heart pump andmethod of controlling a heart pump.

DESCRIPTION OF THE PRIOR ART

The reference in this specification to any prior publication (orinformation derived from it), or to any matter which is known, is not,and should not be taken as an acknowledgment or admission or any form ofsuggestion that the prior publication (or information derived from it)or known matter forms part of the common general knowledge in the fieldof endeavour to which this specification relates.

The use of mechanical device therapy to treat heart failure isincreasing as the general population ages and the number of donor organsfor heart transplantation remains limited. Devices can be used to bridgea patient to heart trans-plant, to recovery, or indeed as a destinationalternative. The latter support strategy requires a device withincreased mechanical durability/lifetime.

A blood pump used for an artificial heart apparatus is required to becompact and have a simple structure and centrifugal pumps or similar aretherefore typically used. In addition, the blood pump must be a pumpthat can continue to rotate without any trouble over a long time. Acontactless rotor using a magnetic bearing apparatus is suitable forsuch a pump apparatus.

Japanese Patent Publication No. 3930834 describes a pump apparatusconfigured by the following: a circular upper permanent magnet isprovided on a top surface of a rotor as a rotating device for acentrifugal pump having a magnetic bearing apparatus; at an upper statorlocated above the rotor is provided suction electromagnetic means thatcooperates with the circular upper permanent magnet to generate asuction force with respect to the rotor in an axial direction; aplurality of lower permanent magnets for rotating the rotor are providedon a bottom surface of the rotor; and at a lower stator located underthe rotor is provided rotation electromagnetic means that cooperateswith the permanent magnets for rotating the rotor to perform rotationalmovement of the rotor.

The human heart has a left ventricle and a right ventricle, and the leftventricle pumps blood through a body, while the right ventricle pumpsblood to lungs for re-oxygenation.

In addition, it is necessary to control the blood pumps so that adischarge amount from the left ventricle blood pump becomes differentthan that from the right ventricle blood pump, but since only one pumpchamber is provided in the pump using the rotating device disclosed inJapanese Patent Publication No. 3930834, two separate pumps, one pumpfor pumping blood through the body and the other pump for pumping bloodto the lungs, are required when applying to the above-described bloodpump, and thus there are problems that a pump apparatus becomes largeand complicated, making it unsuitable for implantation in a subject.

Mechanical durability is dependent on the functionality of the device,in particular, the type of bearings implemented. First generationpulsatile devices necessitate contacting components, which limits theirpredicted mechanical lifetime below three years. The reduced size ofsecond generation non-pulsatile rotary impeller devices has acceleratedthem to the forefront of VAD development.

However, initial techniques for impeller support also imposedsignificant limitations on device lifetime, as they required a shaft,seals and bearings U.S. Pat. No. 4,589,822. Subsequent improvementsresulted in devices that rely on blood immersed pivot support U.S. Pat.No. 5,601,418; however predicted service life is still below five years.

Several techniques have since been developed to improve device lifetime,ranging from complete magnetic suspension U.S. Pat. No. 6,575,717, topassive hydrodynamic suspension U.S. Pat. No. 6,227,797. These thirdgeneration devices eliminate contact wear and reduce the number ofmoving components, potentially increasing lifetime to beyond ten years.These latest generation suspension techniques eliminate any point topoint contact which may also improve the hemolytic performance of thepump.

A number of commercial devices, as well as research devices implementhydrodynamic or magnetic bearing technology. For example, both theVentrassist (Ventracor, Sydney, NSW, AU) and the HVAD (Heatware,Framingham, Massachusetts, USA) incorporate an impeller that iscompletely suspended with hydrodynamic forces and driven with anelectromagnetic motor

U.S. Pat. No. 6,227,797 and U.S. Pat. No. 6,688,861 respectively. TheDuraheart (Terumo, Ann Arbor, Mich., USA) uses an axial magnetic bearingwith a permanent magnetic coupling motor U.S. Pat. No. 6,575,717.Heartmate III (Thoratec, Woburn, Mass., USA) uses a combined radialself-bearing motor U.S. Pat. No. 6,351,048 while the Levacor (formerlyHeartquest, Worldheart, Ottawa, ON, Canada) uses an axial magneticbearing and electromagnetically coupled motor U.S. Pat. No. 6,394,769.

All of the devices mentioned provide left ventricular assistance (LVAD).However, a significant number of patients also require a device forright ventricular assistance (RVAD). The incidence of bi-ventricularfailure is not always initially apparent in heart failure patients, andright ventricular heart failure may develop in up to 40% of patientsreceiving LVAD assistance.

One of the most successful BiVAD techniques used in clinical practiceuses the extracorporeal connection of two Thoratec PVA devices. Smallersecond and third generation rotary systems have also been proposed whichmake use of two separate rotary pumps, such as the combined Coraide andDexaide U.S. Pat. No. 5,890,883 and two Gyro pumps U.S. Pat. No.5,601,418.

However, all currently available bi-ventricular assist systems requirethe use of two devices, with separate controllers, which can introduceleft and right outflow control issues, particularly with the second andthird generation devices. The dual device approach also increasesimplantation size as well as the cost of the therapy.

Single rotary pumps have also been designed to augment the function ofboth ventricles of a failing heart, as described in U.S. Pat. No.5,725,357, U.S. Pat. No. 6,220,832, WO2004098677 and WO2006053384A1.Each of these devices include a double sided impeller that rotates at acommon speed, with each side of the impeller respectively configured forleft and right heart assistance. This effectively introduces an inherentproblem regarding the ability to independently control and thus balancethe outflow from the left and right sides of the device, i.e. anincrease in impeller rotational speed with produce a correspondingincrease in outflow from both cavities.

WO2006053384A1 addressed this issue by introducing the ability toaxially displace the rotating impeller within the cavity so as tosimultaneously alter the relative efficiencies of each side of thedevice. However, this application describes the control method used toachieve this axial displacement as active, thus requiring the use offeedback signals from pressure sensors and the like to actively controland maintain a desired set axial location. This method of control wouldinherently consume excessive amounts of electrical power.

The ability to maintain a balance between the left and right outflow ofa BiVAD system is essential for successful device operation.Haemodynamic parameters that may upset this balance include thebronchial flow, relative changes in systemic and pulmonary vascularresistance, relative changes in left and right ventricularcontractility, pulmonic or systemic congestion, and ventricularcollapse. These conditions infer that a technique for balancing the leftand right VAD hydraulic output is required for long term support.

To operate and control the hydraulic output from each blood pump,parameters such as motor power, speed, differential pressure(inlet-outlet) and flow are required. Whilst determining the motor powerand speed is relatively easy, detecting the remaining parametersconventionally requires additional instrumentation, such as pressuresensors and flow meters. These components increase the possibility ofdevice failure; as such components have limited long term reliability.Furthermore, their addition to the device can induce extra blood contactwith other foreign material, exacerbating the potential for blooddamage.

Previous attempts to regulate the outflow from each device and balancethe left/right outflow requirements have often relied on the use of apressure sensor to detect left atrial pressure (LAP). A feedbackmechanism is then employed to either reduce LVAD speed, or increase RVADspeed, in the presence of reduced LAP. Another technique includes thesurgical introduction of a shunt between the left and right atrium tosafely protect against the potentially disastrous build up of fluid ineither atrium. Alternatively, U.S. Pat. No. 6,527,698 includes a conduitlinking right to left atria through which flow is varied via a variableoccluding valve. However, this technique introduces an additional bloodcontacting conduit, as well as complexities involved with activefeedback control, such as the need for sensors. Furthermore, thissolution can help to balance the fluid distribution but does not providea method for controlling the alteration of device outflow.

As mentioned, the ability to alter the left and right outflow of a BiVADis important, especially in the post operative period when theneuro-humoral auto regulatory mechanisms are least partially ablated byanaesthesia and critical illness.

Many control algorithms exist for the active levitation of magneticbearing systems. While most focus on the maintenance of a centralisedrotor position, an alternative technique exists which focuses on theminimisation of power consumption. The latter controller uses passivelygenerated forces from permanent magnets within the magnetic circuit tocounteract external forces which would otherwise require power from theelectromagnetic coils. This results in movement of the impeller from thecentralised position, until an equilibrium of external and permanentmagnetic forces is reached. Therefore, the power consumption of theactive electromagnetic coils is returned to a minimal state. A number ofrotary blood pump designs implement this form of zero power control.

Masuzawa et. al. (2004) implemented zero power control in a reluctancetype radial magnetic motor bearing to completely suspend the rotor of acentrifugal blood pump. The system places the magnetic materialconcentrically around the rotor, and includes permanent magnets toprovide additional bias flux to the magnetic circuit. These magnets areused by the zero power controller to reduce power consumption whencompared to a central position controller. During operation, theapplication of a radial hydraulic force to the rotor causes atranslation of this rotor in a direction perpendicular to the axis ofrotation, and opposite to the applied force. However, no significanteffect on pump outflow can be observed with this motion, as alteringradial clearance gaps has minimal effect on hydraulic efficiency.(Masuzawa, T., H. Onuma, and Y. Okada, Zero Power Control forMagnetically Suspended Artificial Heart. Jido Seigyo Rengo Koenkai KoenRonbunshu, 2004. 47: p. 322).

U.S. Pat. No. 6,717,311 suspends the rotor of a centrifugal blood pumpin the axial direction with a lorentz type magnetic bearing system. Thissystem again places the magnetic material concentrically around therotor, however the magnetic forces act perpendicularly to the pole face,in the direction parallel to the rotational axis. Additional permanentmagnets, not included in the magnetic bearing circuit, are configured toprovide a counteracting force when an axial hydraulic force isencountered. This counteracting force is effective when allowing thezero power controller to translate the impeller in the same direction asthe applied force. Whilst this motion can be adapted to alter theoutflow of the device, motion in the same direction as the applied forceis undesirable, and will, for example, increase outflow when a decreaseis warranted.

U.S. Pat. No. 6,293,901 uses a lorentz type axial magnetic bearing, alsoconcentrically located around the rotor. Suspension in the radialdirection is achieved using a configuration of repelling permanentmagnets (U.S. Pat. No. 5,928,131,U.S. Pat. No. 6,179,773) configured ina halbach array (U.S. Pat. No. 6,293,901). These magnets are used toachieve zero power control, which relocates the axial position of theimpeller in response to hydraulic axial forces. Since this configurationuses repelling magnets to achieve this, their low stiffness may notprovide sufficient counteraction of force for a given displacement.Although axial relocation is opposite to the direction of the appliedforce, the shrouded configuration as well as the location of theimpeller vanes beneath the impeller does not provide for a reduction inpump outflow in response to the forces generated during instances ofventricular collapse. Therefore, the zero power controller can minimisebearing bower, but not provide flow control based on changing preloadconditions. Furthermore, the mentioned axial gap between the bottomimpeller shroud and casing (0.005 inch) is too small, and impeller bladeheight too large, to produce an appreciable change in hydraulicperformance with the maximum axial translation possible, even if theshroud was semi-open.

US 2007253842 describes a pump includes a housing, a stator supported inthe housing, and a rotor assembly. The rotor assembly includes a rotorsupported in the housing for rotation relative to the stator about anaxis. The rotor assembly also includes a first impeller operativelycoupled to a first axial end of the rotor for rotation with the rotorabout the axis. The rotor assembly further includes a second impelleroperatively coupled to a second axial end of the rotor, opposite thefirst axial end, for rotation with the rotor about the axis. The rotorassembly is movable along the axis relative to the housing to adjusthydraulic performance characteristics of the pump.

SUMMARY OF THE PRESENT INVENTION

The present invention seeks to substantially overcome, or at leastameliorate, one or more disadvantages of existing arrangements.

In a first broad form the present invention seeks to provide a heartpump including:

-   -   a) first and second cavity portions, each cavity portion        including a respective inlet and outlet;    -   b) a connecting tube extending between the first and second        cavity portion;    -   c) an impeller including:        -   i) a first set of vanes mounted on a first rotor in the            first cavity portions;        -   ii) a second set of vanes mounted on a second rotor in the            second cavity portion; and,        -   iii) a shaft connecting the first and second rotors, the            shaft extending through the connecting tube;    -   d) a drive for rotating the impeller; and,    -   e) a magnetic bearing including at least one bearing coil for        controlling an axial position of the impeller, at least one of        the drive and magnetic bearing being provided at least partially        between the first and second cavity portions.

Typically the axial position determines a separation between each set ofvanes and a respective cavity surface, the separation being used tocontrol the fluid flows from the inlets to the outlets.

Typically the drive includes:

-   -   a) a second magnetic material provided in the impeller;    -   b) at least one drive coil that in use generates a magnetic        field that cooperates with the second magnetic material allowing        the impeller to be rotated.

Typically the second magnetic material includes a number ofcircumferentially spaced permanent magnets mounted in the impeller,adjacent magnets having opposing polarities.

Typically the second magnetic material is mounted in the second rotor,and wherein the drive is positioned adjacent the second cavity, thedrive and second magnetic material being configured to result in anattractive force between the drive and the second rotor.

Typically, in use, the at least one bearing coil generates a magneticfield that cooperates with first magnetic material in the impeller,allowing the axial position of the impeller to be controlled.

Typically the first magnetic material is a permanent magnet.

Typically the at least one bearing coil is for generating a magneticfield that is one of complementary to and counter to the first magneticfield generated by the permanent magnet, thereby controlling the netmagnetic field between the bearing and the first magnetic material.

Typically the first magnetic material is mounted in the first rotor andwherein the magnetic bearing is positioned adjacent the first cavity,the magnetic bearing and first magnetic material being configured toresult in an attractive force between the magnetic bearing and the firstrotor.

Typically the impeller includes:

-   -   a) a second magnetic material provided on the second rotor for        cooperating with the drive to allow rotation of the impeller;        and,    -   b) a first magnetic material provided on the first rotor for        cooperating with the magnetic bearing to allow the axial        position of the impeller to be controlled.

Typically:

-   -   a) the magnetic bearing is positioned adjacent the first cavity,        the magnetic bearing and first rotor being configured to result        in a first attractive force between the magnetic bearing and the        second rotor; and,    -   b) the drive is positioned adjacent the second cavity, the drive        and second rotor being configured to result in a second        attractive force between the drive and the second rotor and        wherein the first and second attractive forces are approximately        balanced when the impeller is positioned at an approximately        axially central position during normal circulatory conditions.

Typically the heart pump includes a controller for:

-   -   a) determining movement of the impeller in a first axial        direction;    -   b) causing the magnetic bearing to move the impeller in a second        axial direction opposite the first axial direction;    -   c) determining an indicator indicative of the power used by the        magnetic bearing; and,    -   d) causing the magnetic bearing to control the axial position of        the impeller in accordance with the indicator to thereby control        a fluid flow between the inlets and the outlets.

Typically the controller is for:

-   -   a) comparing the indicator to a threshold; and,    -   b) causing the magnetic bearing to stop movement of the impeller        in the second axial direction depending on the results of the        comparison.

Typically the controller is for minimizing the power used by themagnetic bearing.

Typically the controller is for:

-   -   a) comparing an axial position of the impeller to position        limits; and,    -   b) controlling the magnetic bearing to maintain the axial        position of the impeller within the position limits.

Typically the controller is for:

-   -   a) determining a pressure change within at least part of a        cavity; and,    -   b) controlling the axial position of the impeller in response to        the pressure change.

Typically the controller is for determining the pressure change bydetecting axial movement of the impeller.

Typically the controller is for:

-   -   a) detecting movement of the impeller caused by a change in        fluid pressure within at least one of the cavity portions; and,    -   b) causing the magnetic bearing to control the axial position of        the impeller to thereby change a fluid flow from the inlet to        the outlet for at least one of the cavity portions.

Typically the controller is for, at least one of:

-   -   a) causing the magnetic bearing to reduce the separation between        the vanes and the cavity surface to thereby increase the flow of        fluid from the inlet to the outlet; and,    -   b) causing the magnetic bearing to increase the separation        between the vanes and the cavity surface to thereby decrease the        flow of fluid from the inlet to the outlet.

Typically the controller is for:

-   -   a) detecting movement of the impeller caused by a change in        relative fluid pressures in the cavity portions; and,    -   b) causing the magnetic bearing to control the axial position of        the impeller to thereby alter the relative flow of fluid from        the inlets to the outlets.

Typically the controller is for:

-   -   a) determining axial movement of the impeller away from a normal        balance position;    -   b) causing the magnetic bearing to move the impeller towards the        normal position;    -   c) monitoring the power used by the magnetic bearing;    -   d) determining a new balance position in accordance with the        power used by the magnetic bearing; and,    -   e) causing the magnetic bearing to move the impeller to the new        balance position.

Typically the normal balance position is used to maintain required fluidflows from each inlet to each outlet.

Typically the new balance position is offset from the normal balanceposition.

Typically the new balance position is used to adjust relative fluidflows between the inlets and the outlets.

Typically the indicator is determined using an indication of anelectrical current used by the magnetic bearing.

Typically the controller is for determining a rate of change of currentused by the magnetic bearing to cause axial movement of the impeller.

Typically the controller is for:

-   -   a) determining movement of the impeller in a first axial        direction;    -   b) controlling the magnetic bearing to move the impeller in a        second axial direction opposite the first axial direction until        at least one of:        -   i) the power used by the magnetic bearing falls below a            predetermined amount; and,        -   ii) the axial position of the impeller reaches a position            limit.

Typically, the processing system includes:

-   -   a) a memory for storing instructions; and,    -   b) a processor that executes the instructions, thereby causing        the processor to:        -   i) determine movement of the impeller in the first axial            direction;        -   ii) generate a signal for causing the magnetic bearing to            move the impeller in the second axial direction;        -   iii) determine an indicator indicative of the power used by            the magnetic bearing; and,        -   iv) generate a signal for causing the magnetic bearing to            control the axial position of the impeller in accordance            with the indicator to thereby control a fluid flow between            the inlet and the outlet.

Typically the heat pump includes a controller for:

-   -   a) determining movement of an impeller from a balance position        within a cavity, the cavity including at least one inlet and at        least one outlet, and the impeller including vanes for urging        fluid from the inlet to the outlet;    -   b) causing a magnetic bearing to move the impeller to a new        balance position based on an indication of power used by the        magnetic bearing, the magnetic bearing including at least one        coil for controlling an axial position of the impeller within        the cavity, and the new balance position being used to control        fluid flow from the inlet to the outlet.

Typically the heat pump includes a controller for controlling an axialposition of an impeller within a cavity, a cavity including a firstcavity portion having a first inlet and a first outlet and a secondcavity portion having a second inlet and a second outlet, and theimpeller including first and second sets of vanes, each set of vanesbeing for urging fluid from a respective inlet to a respective outlet,the controller controlling the axial position such that if the relativepressure in the first cavity increases relative to the second cavity,the impeller is positioned in the first cavity thereby increase therelative fluid flows from the first outlet relative to the secondoutlet.

Typically the heat pump includes a controller for controlling an axialposition of an impeller within a cavity, a cavity including an inlet andan outlet, and the impeller including vanes for urging fluid from theinlet to the outlet, the controller controlling the axial position suchthat if the pressure in the cavity increases, the impeller is moves awayfrom the inlet, thereby reducing an outlet flow pressure.

In a second broad form the present invention seeks to provide a methodof controlling a heart pump, the heart pump including:

-   -   a) first and second cavity portions, each cavity portion        including a respective inlet and outlet;    -   b) a connecting tube extending between the first and second        cavity portion;    -   c) an impeller including:        -   i) a first set of vanes mounted on a first rotor in the            first cavity portions;        -   ii) a second set of vanes mounted on a second rotor in the            second cavity portion; and,        -   iii) a shaft connecting the first and second rotors, the            shaft extending through the connecting tube;    -   d) a drive for rotating the impeller; and,    -   e) a magnetic bearing including at least one bearing coil for        controlling an axial position of the impeller, at least one of        the drive and magnetic bearing being provided at least partially        between the first and second cavity portions, the method        including:        -   i) determining movement of the impeller in a first axial            direction;        -   ii) causing the magnetic bearing to move the impeller in a            second axial direction opposite the first axial direction;        -   iii) determining an indicator indicative of the power used            by the magnetic bearing; and,        -   iv) causing the magnetic bearing to control the axial            position of the impeller in accordance with the indicator to            thereby control a fluid flow between the inlets and the            outlets.

In a third broad form the present invention seeks to provide a pumpapparatus comprising:

-   -   a) a stator;    -   b) an upper rotor arranged spaced apart above the stator;    -   c) a lower rotor arranged spaced apart under the stator;    -   d) connecting means provided rotatably and vertically movable at        the stator to connect the upper rotor and the lower rotor;    -   e) first electromagnetic means that is provided on one surface        of the stator opposed to either one of permanent magnets, which        are provided on a bottom surface of the upper rotor and a top        surface of the lower rotor, cooperates with the permanent        magnet, and that generates an acting force with respect to the        rotor in an axial direction to thereby levitate the rotor;    -   f) second electromagnetic means that is provided on the other        surface of the stator opposed to the other permanent magnet, and        that cooperates with the permanent magnet to thereby        rotationally drive the rotor;    -   g) first pumping means at which a first impeller provided on a        top surface of the upper rotor rotates in a first pump chamber;        and    -   h) second pumping means at which a second impeller provided on a        bottom surface of the lower rotor rotates in a second pump        chamber.

Typically the connecting means is composed of an axis of rotationinserted rotatably and vertically movable into a centre through-hole ofthe stator.

The pump apparatus may comprise a control section that moves theconnecting means forward and backward in an axial direction and that canadjust a gap between the first impeller and an opposed surface of thefirst pump chamber, and a gap between the second impeller and the pumpchamber by coordinating them.

It will be appreciated that the broad forms of the invention may be usedindividually or in combination.

BRIEF DESCRIPTION OF THE DRAWINGS

An example of the present invention will now be described with referenceto the accompanying drawings, in which:—

FIG. 1 is a sectional view of an example of a pump apparatus;

FIG. 2 is a perspective view of an example of the first and secondrotors of the pump apparatus of FIG. 1;

FIG. 3 is a block diagram of an example of a controller for the pumpapparatus of FIG. 1;

FIG. 4 is a sectional view of an example of a pump apparatus of FIG. 1with the rotor in a first axial position;

FIG. 5 is a sectional view of an example of a pump apparatus of FIG. 1with the rotor in a second axial position;

FIG. 6A shows a schematic cross sectional view of a second example of aheart pump;

FIG. 6B shows a schematic cross sectional view of a third example of aheart pump;

FIG. 7 is a flow chart of an example of a method of controlling aposition of an impeller in a heart pump;

FIG. 8 is a schematic diagram of an example of a controller;

FIG. 9 is a flow chart of a second example of a method for controlling aposition of an impeller in a heart pump;

FIG. 10A is a schematic perspective exploded view of an example of driveand magnetic bearing system for a heart pump;

FIG. 10B is a schematic side view of the drive and magnetic bearingsystem of FIG. 10A;

FIG. 10C is a schematic plan view of the drive system of FIG. 10A;

FIG. 10D is a schematic plan view of the magnetic bearing system of FIG.10A;

FIG. 10E is a schematic plan view of the rotor of the drive system ofFIG. 10A;

FIG. 10F is a schematic plan view of the rotor of the magnetic bearingsystem of FIG. 10A;

FIG. 11A is a schematic side illustrating the operation principle ofaxial position control by the magnetic bearing;

FIG. 11B is a schematic side illustrating the operation principle oftilt control by the magnetic bearing;

FIGS. 12A to 12C are schematic side views illustrating the manner inwhich the impeller position is relocated with zero power control;

FIGS. 13A to 13C are schematic side views illustrating the manner inwhich a single sided impeller position is relocated to influence deviceoutflow;

FIGS. 14A to 14C are schematic side views illustrating the manner inwhich a dual sided impeller position is relocated to influence deviceoutflow;

FIG. 15 is a graph illustrating the relative hydraulic performance ofthe left and right flows in an example BiVAD;

FIGS. 16A to 16H are schematic diagrams of an example of a BiVAD;

FIGS. 17A to 17D are schematic diagrams illustrating the resultant axialforce development on the impeller of a BiVAD during a variety of commonconditions;

FIGS. 18A to 18C are schematic diagrams illustrating the adaptation of aBiVAD to alterations in relative vascular resistance;

FIGS. 19A to 19C are schematic diagrams illustrating the adaptation of aBiVAD to heart chamber collapse;

FIGS. 20A to 20D are schematic diagrams illustrating the resultant axialforce development on the impeller of a VAD during a variety of commonconditions;

FIGS. 21A to 21C are schematic diagrams illustrating the adaptation of aVAD to alterations in relative vascular resistance; and,

FIGS. 22A to 22C are schematic diagrams illustrating the adaptation of aVAD to heart chamber collapse.

FIG. 23A is a graph of example left and right atrial pressures (LAP,RAP) for an example heart pump in a circulatory loop simulating a numberof circulatory conditions;

FIG. 23B is a graph of example systemic and pulmonary resistances (SVR,PVR) for an example heart pump in a circulatory loop simulating a numberof circulatory conditions;

FIG. 23C is a graph of example impeller axial positions for an exampleheart pump in a circulatory loop simulating a number of circulatoryconditions;

FIG. 23D is a graph of example magnetic bearing power (MB power) usagefor an example heart pump in a circulatory loop simulating a number ofcirculatory conditions; and,

FIG. 23E is a graph of example aortic and pulmonary pressures (AoP, PAP)for an example heart pump in a circulatory loop simulating a number ofconditions.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

A first example of a pump apparatus will now be described with referenceto FIGS. 1 to 5. In particular, FIG. 1 is a sectional view of a pumpapparatus 1, and FIG. 2 is a perspective view of first and second rotorsof the pump apparatus 1.

In this example, a first pump 2 is arranged in a first part 1 a of thepump apparatus 1, and a second pump 3 is arranged in a second part 1 bof the pump apparatus 1. In this example, at least part of a magneticbearing and/or a drive, in this example a stator 4, is provided betweenthe first and the second pumps 2, 3.

The pump apparatus includes an impeller 5, having a first rotor 5 a anda second rotor 5 b arranged in the first and second parts 1 a, 1 b, withthe stator 4 being provided therebetween. The first and second rotors 5a, 5 b are connected via a shaft 6 rotatably mounted within the pumpapparatus to allow rotation about, and axial movement along an axis 6 ain the connecting tube 4 a, which in this example extends through thestator 4.

Thus, it will be appreciated that in this example, the connecting tube 4a is provided as a centre-hole through the stator 4, so that the stator4 is provided radially outwardly of the connecting tube 4 a. However,this is not essential and other arrangements can be used. For example,the connecting tube could be annular in shape, with the shaft being ahollow cylinder or a plurality of rods extending through the annularconnecting tube.

The first rotor 5 a is disk-shaped, and supports a first impeller 2 acomprising a plurality of vanes that are provided on a first surfacethereof. A first magnetic material in the form of a permanent magnet 7 ais provided on a second surface of the first rotor 5 a, facing thestator 4. The second rotor 5 b is also disk-shaped, and supports asecond impeller 3 a composed of a plurality of vanes that are providedon a first surface thereof, with a second magnetic material, in the formof a plurality of permanent magnets 7 b radially arranged on a secondsurface thereof, facing the stator 4. The permanent magnets 7 btypically include a number of circumferentially spaced permanent magnetsmounted in the impeller 5, adjacent magnets having opposing polarities,however, other suitable arrangements may be used, such as a Halbacharray.

The stator 4 is composed of a doughnut-shaped body 4 b, and firstelectromagnetic means 8 a composed of, for example, four electromagnets,provided on a first surface of the body 4 b, facing the first rotor 5 a.The first electromagnetic means 8 a generates a magnetic field thatcooperates with permanent magnet 7 a, allowing the axial position of theimpeller 5 to be controlled.

Thus, the first electromagnetic means 8 a and the first magneticmaterial form a magnetic bearing including at least one bearing coil forcontrolling an axial position of the impeller 5. In one example, the atleast one bearing coil generates a magnetic field that is one ofcomplementary to and counter to the magnetic field generated by thepermanent magnet 7 a, thereby controlling the net magnetic field betweenthe bearing and the first magnetic material, allowing the axial positionof the first rotor 5 a, and hence the impeller 5 to be altered.

In addition, second electromagnetic means 8 b is provided on a secondsurface of the body 4 b facing the second rotor 5 b that is, forexample, composed of twelve three-phase electromagnets for generating arotating magnetic field, whereby the electromagnets and the permanentmagnet 7 b cooperate with each other to thereby rotationally drive thesecond rotor 5 b.

Accordingly, the second electromagnetic means 8 b and magnets 7 b form adrive for rotating the impeller. The drive therefore typically includesa second magnetic material provided in the impeller and at least onedrive coil that in use generates a magnetic field that cooperates withthe second magnetic material allowing the impeller to be rotated.

The first pump 2 includes a first pump chamber (or cavity portion) 2 band the first impeller 2 a. The first impeller 2 a can move axiallywhile rotating with the first rotor 5 a in the first pump chamber 2 b. Afirst inlet port 2 c is provided in a centre of an outer surface wall ofthe first pump chamber 2 b, with an outlet port 2 d being provided on aside wall of the first pump chamber 2 b.

The second pump 3 includes a second pump chamber 3 b and the secondimpeller 3 a. The second impeller 3 a can move axially while rotatingwith the second rotor 5 b in the second pump chamber 3 b. A second inletport 3 c is provided in a centre of an outer surface wall of the secondpump chamber 3 b, with an outlet port 3 d being provided on a side wallof the second pump chamber 3 b.

It will be appreciated that the pump chambers 2 b, 3 b may be providedwith volutes to thereby assist with transfer of the fluid to the outlets2 d, 3 d. The volutes maybe any combination of type spiral/single,split/double or circular/concentric, however the latter circular volutetype is preferred if a journal bearing is used, as will be described inmore detail below, as this configuration produces a stabilising radialhydraulic force for optimal journal bearing functionality.

FIG. 3 is a block diagram of a control system of the pump apparatus 1 ofthe present example.

In FIG. 3, reference numeral 9 a denotes a first sensor that is providedcorresponding to each electromagnet of the first electromagnetic means 8a of the first surface of the stator 4 and that is composed of eddycurrent sensors. The first sensor 9 a detects an axial position of thefirst rotor 5 a and inputs a detected signal into a controller 11through an A/D converter 10. Additionally, inclination of the rotors iscalculated in the controller 11 into which the detected signal is inputfrom the first sensor 9 a as will be explained hereinafter.

In this example, four first sensors 9 a are provided corresponding tothe four electromagnets 8 a, with first detected signals of axialpositions of the first rotor 5 a and the second rotor 5 b being obtainedby calculating an average value of the detected signals from the fourfirst sensors 9 a in the controller 11. Additionally, second detectedsignals of inclination of the first and second rotors 5 a and 5 b areobtained from an average difference of the adjacent first sensors 9 a,and from a distance between the first sensors 9 a.

Note that an example using four first sensors 9 a is explained above,but the inclination can be calculated in a case of the three or morefirst sensors 9 a. Furthermore, in one example, particularly if ahydrodynamic bearing is used, inclination or tilt sensing may not berequired.

Reference numeral 9 b denotes a second sensor that is provided near thesecond rotor 5 b and that detects rotational speeds of the first andsecond rotors 5 a and 5 b. The second sensor 9 b inputs third detectedsignals of the rotational speeds into the controller 11 through the A/Dconverter 10.

The controller 11 implements a PID (Proportional-Integral-Derivative)control operation based on the first and second detected signals,outputs a control signal to the first electromagnetic means 8 a througha drive circuit 12 a, thereby controlling the axial position and theinclination of the first rotor 5 a. Simultaneously, the controllercarries out a PID control operation based on the third detected signals,outputs a control signal to the second electromagnetic means 8 b througha drive circuit 12 b, and controls the rotational speed of the secondrotor 5 b.

Next, usages and operations of the pump apparatus 1 of the presentembodiment will be explained.

In one example, the pump apparatus 1 is used for a blood pump of anartificial heart. Namely, the first pump 2 operates so as to pump bloodthrough the body as a left ventricle blood pump, and additionally, thesecond pump 3 operates so as to pump blood to the lungs as a rightventricle blood pump.

Rotational drive of the above-described pump is performed cooperativelyby the permanent magnet 7 b of the second surface of the second rotor 5b and the second electromagnetic means 8 b of the second surface of thestator 4. The rotational speed of the first rotor 5 a is the same asthat of the second rotor 5 b by virtue of the physical coupling betweenthe rotors 5 a, 5 b, provided by the shaft 6.

In addition, during the rotation of the rotors 5, the firstelectromagnetic means 8 a and the permanent magnet 7 a cooperate witheach other to thereby level the first rotor 5 a and the second rotor 5 bwith an inclination control signal from the controller 11. Additionally,an attractive force between the first electromagnetic means 8 a and thepermanent magnet 7 a is adjusted in accordance with a control signal ofthe axial position from the controller 11. This is performed in order toadjust a flow ratio of the first pump 2 to the second pump 3, byaltering the axial position of the first rotor 5 a and the second rotor5 b.

In particular, a force towards the second pump 3, generated as a sum ofthe attractive force between the first electromagnetic means 8 a and thepermanent magnet 7 a, can be adjusted as described above. Accordinglythis controls the levitation of the rotors 5 a, 5 b while adjusting theaxial positions of the rotors 5 a, 5 b, and hence the impeller 5. Thisis balanced with an opposing force, towards the first pump 2 generatedby an attractive force between the second electromagnet 8 b and thepermanent magnet 7 b.

Thus, the drive is positioned adjacent the second cavity portion, andwherein the drive and second magnetic material are configured to resultin an attractive force between the drive and the second rotor, whilstthe magnetic bearing is positioned adjacent the first cavity, with themagnetic bearing and first magnetic material being configured to resultin an attractive force between the magnetic bearing and the first rotor.In one example, the attractive forces are approximately balanced whenthe impeller is positioned at an approximately axially central positionduring normal circulatory conditions.

When the levitation of the rotors 5 a, 5 b is controlled and then thefirst rotor 5 a moves upwardly, a gap d₁ between the first impeller 2 aand a top surface wall of the first pump chamber 2 b becomes smaller asshown in FIG. 4. At the same time, a gap d₂ between the second impeller3 a and a bottom surface wall of the second pump chamber 3 b becomeslarger. At this time, since the gap d₁ between the first impeller 2 aand the top surface wall of the first pump chamber 2 b is smaller in thefirst pump 2, backflow hardly occurs and an outlet flow from the outletport 2 d increases. Since the gap d₂ between the second impeller 3 a andthe bottom surface wall of the second pump chamber 3 b is large in thesecond pump 3, even if discharge is performed by the second impeller 3a, backflow occurs in the gap d₂ and the outlet flow from the outletport 3 d becomes smaller.

In contrast, when the first rotor 5 a moves downwardly, towards thestator 4, the gap d₁ between the first impeller 2 a and the top surfacewall of the first pump chamber 2 b increases as shown in FIG. 5. The gapd₂ between the second impeller 3 a and the bottom surface wall of thesecond pump chamber 2 b also becomes smaller. At this time, since thegap d₁ between the first impeller 2 a and the top surface wall of thefirst pump chamber 2 b is larger in the first pump 2, even if dischargeis performed by the first impeller 2 a, backflow occurs in the gap d₁and thereby an outlet flow from the outlet port 2 d decreases. Since thegap d₂ between the second impeller 3 a and the bottom surface wall ofthe second pump chamber 3 b is smaller in the second pump 3, backflowhardly occurs and an outlet flow from the outlet port 3 d increases.

Consequently, if a discharge amount of the left ventricle is increased,when the rotor is moved towards the first pump chamber 2 b, thedischarge amount of the left ventricle becomes large, and that of theright ventricle becomes small as described above.

Hence, the positions of the axial movement of the rotors 5 may becontrolled by the control signals from the controller 11 so that a ratioof the discharge amount of the left ventricle to that of the rightventricle becomes a desired one.

Accordingly, in this example, the axial position of the impellerdetermines a separation between each set of vanes and a respectivecavity surface, the separation being used to control the fluid flowsfrom the inlets to the outlets. In one further example, the controllercan therefore control the flows from the inlets to the outlets to alterthe fluid flow to take into account different haemodynamic states. Thiscan be achieved by determining movement of the impeller in a first axialdirection, causing the magnetic bearing to move the impeller in a secondaxial direction opposite the first axial direction, determining anindicator indicative of the power used by the magnetic bearing andcausing the magnetic bearing to control the axial position of theimpeller in accordance with the indicator to thereby control a fluidflow between the inlets and the outlets. An example of such control willbe described in more detail below.

Note that in the present example, an example is explained that thepermanent magnet 7 a is engaged with the bottom surface of the firstrotor 5 a, but the first rotor 5 a may be formed of a magnetic materialinstead of engaging the permanent magnet 7 a thereto.

In addition, in the present embodiment, an example is explained that thefirst rotor 5 a and the second rotor 5 b are connected to be fixed tothe shaft 6 that passes through the centre through-hole of the stator 4,but they may be connected to be fixed with a plurality of rods locatedoutside and near the side surface of the stator 4 and capable ofvertical movement.

The pump apparatus 1 described above is a compact, simple structure, andis also controlled easily, so that the structure is the most suitablefor a blood pump of an artificial heart. The apparatus can also be usedfor a blood pump of an artificial heart or for a pump that discharges afluid by simultaneously adjusting a ratio of discharge amounts from thepump chambers 2 b, 3 b. Accordingly, in one example, a pump apparatus isprovided that can easily control each discharge amount of a leftventricle blood pump and a right ventricle blood pump as a blood pump ofan artificial heart.

The above described arrangement provides a number of advantageouseffects. For example, the pump can provide a function of the leftventricle blood pump and a function of the right ventricle blood pump inone pump apparatus, which is compact and lightweight and thereby theequipment can be simplified, and which can be driven without any troubleover a long time by magnetic levitation means, and further in whichdischarge amounts of the right and left blood pumps can be easilycontrolled simultaneously.

In one example, a first rotor 5 a above a stator 4 and a second rotor 5b thereunder are connected to each other with an axis of rotation 6 ainserted into a centre through-hole 4 a of the stator 4, and an actingforce in an axial direction is generated by first electromagnetic means8 a of a first surface of the stator 4 with respect to a permanentmagnet 7 a of a second surface of the first rotor 5 a, while secondelectromagnetic means 8 b of a second surface of the stator 4 operateson a permanent magnet 7 b of a first surface of the lower rotor 5 b torotationally drive the second rotor 5 b, whereby a first portion of thepump apparatus operates as first pumping means 2 by first impeller vanes2 a on a first surface of the first rotor 5 a, while a second portion ofthe pump apparatus operates as second pumping means 3 by second impellervanes 3 a of a second surface of the second rotor 5 b.

A reference numerals list is set out below, with alternative terminologythat can be used interchangeably with the terminology above beinglisted:

-   1—Heart Pump or Pump Apparatus-   1 a—First Part-   1 b—Second Part-   2—First Pump or First Pumping Means-   2 a—First Impeller-   2 b—First Cavity Portion or First Pump Chamber-   3—Second Pump-   3 a—Second Impeller-   3 b—Second Cavity Portion or Second Pump Chamber-   4—Stator-   4 a—Connecting Tube or Center hole-   5 a—First rotor or Upper Rotor-   5 b—Second rotor or Lower Rotor-   6—Shaft-   6 a—Axis of Rotation-   7 a—First Magnetic Material or First Permanent magnet-   7 b—Second Magnetic Material or Second Permanent magnet-   8 a—First Electromagnetic Means-   8 b—Second Electromagnetic Means-   11—Controller or Control Section

A second example of a heart pump will now be described with reference toFIG. 6A.

In this example, the heart pump 100A includes a housing 110 defining acavity 120, containing an impeller 130A. The impeller 130A effectivelydivides the cavity 120 into first and second cavity portions 121, 122A.The housing 110 includes first and second inlets 141, 142 andcorresponding first and second outlets 151, 152, which are in fluidcommunication with the first and second cavity portions 121, 122A,respectively.

The impeller 130A includes first and second sets of vanes 131, 132, suchthat rotation of the impeller 130A about a rotation axis 160 urges fluidfrom the inlets 141, 142 to the corresponding outlets 151, 152. In use,rotation of the impeller 130A is achieved using a drive, such as amagnetic drive 170. The magnetic drive 170 typically includes at leastone coil positioned at a first end of the housing 110 adjacent the firstcavity 121. In use, the coil generates a magnetic field that cooperateswith magnetic material in the impeller 130A, allowing the impeller to berotated. This tends to lead to an attractive force between the drive 170and the impeller 130A that urges the impeller 130A in an axial directiontowards the first cavity 121.

In use, a relative physical separation between the set of vanes 131, 132and the corresponding cavity surfaces 123, 124A controls the relativeefficiency of the vanes 131, 132 and hence the relative flows betweenthe inlets 141, 142 and the corresponding outlets 151, 152. The positionof the impeller 130A in an axial direction is typically controlled usinga magnetic bearing 180. The magnetic bearing 180 typically includes atleast one coil positioned at a second end of the housing 110 adjacentthe second cavity 122A. In use, the coil generates a magnetic field thatcooperates with magnetic material also in the magnetic bearing stator,which interacts with ferrous material within the impeller 130A, allowingthe axial position of the impeller 130A to be controlled. This tends tolead to an attractive force between the magnetic bearing 180 and theimpeller 130A, that urges the impeller 130A in an axial directiontowards the second cavity 122A.

Although the drive 170 and bearing 180 are positioned adjacent the firstand second cavity portions 121, 122A respectively, this is not essentialand the positioning could be reversed, with the bearing 180 beingpositioned adjacent the first cavity portion 121 and the drive 170positioned against the second cavity portion 122A. However, in general,as the drive needs to provide a rotational torque to the impeller 130A,this is easiest if the drive couples to the impeller 130A in the firstcavity portion as the impeller 130A has a greater diameter at thispoint, which in turn maximises the torque of the drive 170.

The drive 170 and the bearing 180 are typically coupled to a controller190, allowing operation of the heart pump 100 to be controlled. Thecontroller is also typically coupled to a sensor, an example of which isdescribed in more detail below, allowing the position of the impeller tobe determined.

A third example of a heart pump will now be described with reference toFIG. 6B. In this example, similar reference numerals are used todesignate similar features, and these will not therefore be described inany detail.

In this example, the heart pump 100B includes a modified second cavity112B, having a surface 124B that extends across the housing 110, wherethe inlet is provided in the example of FIG. 6A. Accordingly, in thisexample, the heart pump 110 does include a second inlet or a secondoutlet. Furthermore, impeller 130B includes only a single set of vanes131, positioned in the cavity 121, and includes an aperture 135extending through the impeller 130B, for allowing blood to flow from thesecond cavity 122B to the first cavity 121, to thereby preventstagnation between the impeller 130B and the second cavity surface 124B.

In use, the heart pumps 100A, 100B can be coupled to a subject tosupplement the pumping action of one or both of the left and rightventricles of the heart.

For example, the heart pump 100A of FIG. 6A can be coupled to both thepulmonary and systemic circulatory systems, allowing the pump to operateas a BiVAD (Bi-Ventricular Assist Device), in which the pump supplementsthe pumping action of both the left and right ventricles of the heart.In this instance, the left ventricle and the right atrium are coupled tothe first and second inlets 141, 142 respectively, whilst the first andsecond outlets 151, 152 and provide outflow to the aorta and thepulmonary artery, respectively.

In use, the heart pump 100A is arranged so that with the impeller 130Apositioned at an approximately axially central point within the cavity120, generally referred to as a nominal balance point, the pumpingaction provided by each set of vanes 131, 132 equates to the pumpingaction required by each of the left and right ventricles respectively.This can be achieved by selection of suitable dimensions, such as thelength, height and shape of the respective vanes, and in generalachieves a flow of approximately 5 L/min at each outlet 151, 152.

When the circulatory system is functioning correctly, the pressurewithin the first cavity portion 121 will be greater than the pressure inthe second cavity portion 122A, approximately 100 mmHg as opposed to 20mmHg. In this instance, blood flow between the first and second cavities121, 122A is substantially prevented due to the presence of the impeller130A. Depending on impeller geometry, this normal pressure differentialmay lead to a force on the impeller 130A, for example acting towards thesecond cavity portion 122A.

In one example, the heart pump is naturally balanced, so that any suchforces on the impeller 130A including forces resulting from the pressuredifferential and the attractive forces arise caused by magnetic couplingbetween the impeller 130A and the drive 170, as well as between theimpeller 130A and the bearing 180, are approximately equal when theimpeller 130A is provided at a balance point.

As will be described in more detail below, the position of the balancepoint within the cavity 120 is controlled, and is typically positionedat an axial centre of the cavity 120 when the circulatory system isfunctioning correctly. This can be achieved by selection of suitablemagnetic properties for the impeller 130A, the drive 170, as well as thebearing 180. In such a situation, the additional force that is requiredto be exerted by the magnetic bearing 180 to maintain the impeller 130Aat the balance point is minimal, which is sometimes referred to as a“zero power configuration”. In this regard, the term zero is understoodto not necessarily be zero, but rather means that the power required isless than if such balancing were not present.

In this example, if there is a change in the relative pressures of thefirst and second cavity portions 121, 122A caused by an increase inpressure within the pulmonary circulatory system, and/or a decrease inpressure within the systemic circulatory system, then this will lead toa modified pressure differential. The modified pressure differentialresults in a net force on the impeller 130A, causing movement of theimpeller 130A away from the balance point towards the first cavityportion 121, thereby reducing separation between the first set of vanes131 and the first cavity surface 123. This increases the efficiency ofthe pumping effect in the systemic system, and decreases the pumpingeffect in the pulmonary system. This will reduce flow into the pulmonarycirculatory system, whilst increasing the flow into the systemiccirculatory system, which in turn will exacerbate the flow balancingproblems in the pulmonary and/or systemic circulatory system.

Similarly, a decrease in pressure within the pulmonary circulatorysystem, and/or an increase in pressure within the systemic circulatorysystem, will also result in a net force causing movement of the impeller130A towards the second cavity portion 122A. This reduces separationbetween the second set of vanes 132 and the second cavity surface 124A,thereby increasing the efficiency of the pumping effect in the pulmonarysystem, and decreasing the pumping effect in the systemic system, thusincreasing the flow balancing problems in the pulmonary or systemiccirculatory systems.

To address this situation, a control process is implemented, typicallyby the controller 190, which allows the position of the impeller 130Awithin the cavity 120 to be controlled, as will be described in moredetail below.

It will be appreciated that the heart pump 1 described above withrespect to FIGS. 1 to 5 can function in a similar manner, and this willnot therefore be described in further detail.

However, it is also notable that differences exist between the heartpumps 1, 100A. In particular, in heart pump 1, the cavity portions 2, 3are separated by a connecting tube 4 a, whereas in FIG. 6A, the cavityportions 121, 122A are separated only by the impeller 130A. As a result,the minimum cross sectional area and the distance of the path alongwhich fluid has to travel if it is to flow from one cavity portion toanother is greatly increased in the heart pump 1 of FIGS. 1 to 5 thancompared to the heart pump 100A. This helps reduce leakage between theheart pump cavity portions 2, 3, which can be desirable in somesituations.

Leakage can additionally be controlled through appropriate selection ofthe relative diameters of the shaft 6 and the connecting tube 4 a, whichact to alter the minimal cross sectional area along the flow path, aswell as through other factors, such as the presence of flow elements,such as ridges or troughs on the shaft or connecting tube surfaces.

In the example of FIGS. 1 to 5, the first and second electromagneticmeans 8 a, 8 b act as a bearing and drive respectively. Accordingly, itwill be appreciated that in this example, the drive 8 b couples to thesecond rotor 5 b, which has a diameter smaller than that of the firstrotor 5 a. Typically the diameters are 30-40 mm for the second rotor 5b, and 50-60 mm for the first rotor 5 a.

In any event, coupling of the drive to a smaller part of the impeller 5is the reverse of the preferred implementation of the heart pump 100A,in which the drive 170 couples to the larger diameter portion of theimpeller 130A. In one example, this is as a result of a reduceddifferential diameter between the first and second rotors 5 a, 5 b,thereby reducing the torque benefit that could be obtained coupling thedrive to the larger diameter portion of the impeller 5. However, it willbe appreciated that any suitable arrangement could be used, so that thedrive could couple to the larger rotor 5 a.

In the example of the heart pump 100B of FIG. 6B this can be coupled toeither the pulmonary or systemic circulatory systems, allowing the pumpto operate as a VAD (Ventricular Assist Device), in which the pumpsupplements the pumping action of both either the left or rightventricles of the heart.

In this example, the heart pump 100B is arranged so that with theimpeller 130B positioned at an approximately axially central pointwithin the cavity 120, generally referred to as a nominal balance point,the pumping action equates to the pumping action required the respectiveventricle. This can be achieved by selection of suitable dimensions,such as the length, height and shape of the respective vanes, and ingeneral achieves a flow of approximately 5 L/min at the outlet.

In this example, the fluid pressure within the cavities 121, 122B willresult in a pressure differential on the impeller 130B, depending on theconfiguration of the impeller 130B, as will be described in more detailbelow. The pump 100B is again naturally balanced, so that any forces onthe impeller 130B, including forces resulting from the pressuredifferential and the attractive forces arise caused by magnetic couplingbetween the impeller 130B and the drive 170, as well as between theimpeller 130B and the bearing 180, are approximately equal when theimpeller 130B is provided at a balance point. Again, the balance pointwithin the cavity 120 is typically positioned at an axial centre of thecavity 120 when the circulatory system is functioning correctly, so a“zero power configuration”, is implemented.

In this example, if there is a change in the pressure within thecavities 121, 122B, this will result in a modified pressure differentialon the impeller, causing movement of the impeller 130B away from thebalance point. The direction in which this occurs depends on thepressure differential generated across the impeller 130B, which againcould exacerbate flow problems.

An example of a control process will now be described with reference toFIG. 7. This control process is equally applicable to the heart pump ofFIGS. 1 to 5, as well as to the second and third example heart pumps100A, 100B, of FIGS. 6A and 6B. Accordingly, the following descriptionwill focus generally on a heart pump 100, having a impeller 130, withthe respective heart pump 100A, 100B being identified only when this hasan impact on the process, and with the heart pump 1 being assumed tofunction in a manner substantially similar to that described withrespect to the heart pump 100A.

In this example, at step 200, the process includes determining movementof the impeller 130 in a first axial direction in accordance withsignals from a sensor 195. The sensor can be adapted to detect movementof the impeller 130 in any suitable manner. For example, this could beachieved through the use of pressure sensors capable of detectingchanges in the relative pressures within the first and second cavityportions 121, 122, or within the systemic and pulmonary circulatorysystems, which would in turn lead to movement of the impeller 130.Alternatively, this could be achieved by detecting a separation betweenthe impeller 130 and a sensor 195, such as a suitable position sensor,as will be described in more detail below.

It will be appreciated that the first direction could be either towardsthe first or second cavity portions 121, 122, depending on the nature ofthe variation from the normal pressure differential, and in particularwhether this is caused by an increase and/or decrease in either one orboth of the systemic and pulmonary systems.

At step 210, the magnetic bearing 180 is used to move the impeller 130in a second axial direction opposite the first axial direction. Thus,for example, this can involve increasing or decreasing the force appliedby the magnetic bearing 180, thereby allowing the impeller 130 to bemoved against the force caused by the change to the normal pressuredifferential between the two cavity portions 121, 122.

At step 220, an indicator indicative of the power used by the magneticbearing 180 is determined. This may be achieved in any suitable manner,such as by monitoring the current drawn by the magnetic bearing 180.This could include, for example, monitoring the current used to maintainthe impeller 130 at a constant axial position, or alternatively caninvolve monitoring the change in current drawn by the magnetic bearing180 as the impeller 130 is moved in the second axial direction.

At step 230, the indicator is used to control the axial position of theimpeller 130. This is typically performed so as to minimise the currentdrawn by the magnetic bearing 180, within any required constraints,thereby maintaining a zero power configuration, even in the presence ofa change in the normal pressure differential. This situation results ina new balance position that is offset from the axial central position.In particular, in the above described configuration, as there areattractive forces between the impeller 130 and both the drive 170 andthe magnetic bearing 180, the new balance position will be offset fromthe axial central position in the second direction.

Thus, if the impeller 130 is urged towards the second cavity 122, by anincrease in pressure within the first cavity 121, then the new balanceposition will be located offset from the axially central positiontowards the first cavity 121. In this position, the attractive forcebetween the drive 170 and the impeller 130 is increased, whilst theattractive force between the magnetic bearing 180 and the impeller 130is decreased. This results in a greater net force towards the firstcavity 121, balancing the increased force towards the second cavity 122that is caused by the change in pressure differential.

In the example of the heart pump 100A, with the new balance positionoffset towards the first cavity 121, the reduced separation between thefirst set of vanes 131 and the first cavity surface 123 increases thepumping effect within the first cavity 121, thereby increasing the flowfrom the first outlet 151. Similarly there is a decrease in the pumpingeffect in the second cavity 122A, resulting in a reduction in flow fromthe second outlet 152. Accordingly, this operates to restore the normalflow balance of the circulatory system. It will be appreciated that oncethe normal pressure differential is restored, this will result in a netforce towards the first cavity 121.

Applying the above described process will then return the impeller 130Ato the normal balance position axially centred within the cavity 120.

Thus, for the heart pump 100A, the above described process operates tomaintain zero power control by adjusting the position of the impeller130A within the cavity 120 when a change in the normal pressuredifferential arises between the circulatory systems, allowing theimpeller 130A to be provided at a new balance position. Furthermore, inone example, the inherent attractive forces between the impeller 130Aand the drive 170 and magnetic bearing 180 result in a new balanceposition offset from the axial centre of the cavity 120, therebycontrolling relative flows between the inlets 141, 142 and thecorresponding outlets 151, 152, which in turn can be used to compensatefor changes in vascular resistance in the circulatory systems.

It will therefore be appreciated that this control process allows theheart pump 100A to implement relative flow control whilst maintaining azero power configuration.

In the example of the heart pump 100B, an increase in pressure withinthe cavities 121, 122B results in a pressure differential across theimpeller 130B that moves the impeller 130B towards the first cavity 121.In this instance, the new balance position will therefore be offsettowards the second cavity 122B, with the increase separation between thefirst set of vanes 131 and the first cavity surface 123 leading to adecrease in the pumping effect within the first cavity 121. This reducesthe flow pressure, thereby counteracting the increased flow pressurewithin the first and second cavities 121, 122B.

Thus, for the heart pump 100B, the above described process operates tomaintain zero power control by adjusting the position of the impeller130B within the cavity 120 when a change in the normal pressuredifferential across the impeller arises due to a change in pressurewithin the respective circulatory system to which the pump 100B isattached. Unlike the BiVAD application for the pump 100A, correcting forthe pressure differential is more important than correcting relativeflows between the circulatory systems, as relative flow will beinfluenced by the ventricle not attached to the pump 100B. Thus, in theVAD example for the pump 100B, the inherent attractive forces betweenthe impeller 130B and the drive 170 and magnetic bearing 180 result in anew balance position offset from the axial centre of the cavity 120,thereby counteracting any pressure changes, which in turn can be used tocompensate for changes in vascular pressure in the circulatory systems.

It will therefore be appreciated that this control process allows theheart pump 100B to implement pressure control whilst maintaining a zeropower configuration. In this example, if flow control is required, thiscan be achieved by adjusting the rotation speed of the impeller 130B. Itwill be appreciated that this is less complex than attempting to alterthe rotation rate of the impeller 130A in the heart pump 100A, whichcould result in an undesirable flow differential arising due thedifferent configuration of the first and second sets of vanes 131, 132.

Whilst the above described process may be achieved in any suitablemanner, in one example, this is achieved by the controller 190.Accordingly, the controller 190 is typically adapted to control thedrive 170 to cause rotation of the impeller 130. In addition to this,the controller 190 will monitor the power drawn by the magnetic bearing180, and use this to control the magnetic bearing 180, and hence theaxial position of the impeller 130, as will be described in more detailbelow.

It will be appreciated from this that any suitable form of controllermay be used, and an example controller will now be described in moredetail with respect to FIG. 8.

In this example the controller 190 includes a processor 300, a memory301, an optional input/output device (I/O device) 302, such as inputbuttons, a key pad, display or the like, or an optional externalinterface 303, allowing the controller 190 to receive signals from thesensor 195, and provide control signals to the drive 170 and themagnetic bearing 180. It will therefore be appreciated that thecontroller 190 may be in the form of a suitably programmed processingsystem, such as a computer, laptop, palm top, PDA, or alternatively maybe specialised hardware, a programmable logic controller, fieldprogrammable gate array (FPGA) or the like.

In one example, the controller 190 is formed from custommicro-electronics, allowing the controller 190 to be physicallyimplanted together with the heart pump, in a subject. Alternatively, thecontroller 190 could be used to control the heart pump 100 via wirelessconnections, or the like.

In use, the processor 300 executes instructions, typically stored in thememory 301, allowing the processor 300 to perform the control processesdescribed herein. In particular, the processor 300 receives signals fromthe sensors 195 to allow an impeller position to be determined. Theprocessor 300 then determines if modification of the operation of thebearing 180, and/or the drive 170 is required, and if so generatesappropriate signals that are applied to the bearing and/or drive asrequired.

It will be appreciated that the controller 190 could also be implementedin the heart pump 1 of FIGS. 1 to 5 by replacing the controller 11.

An example of a process for using the controller 190 to control theposition of a heart pump impeller will now be described with referenceto FIG. 9. For the purpose of example, this will be described withrespect to the heart pump 100A of FIG. 6A, although it will beappreciated that similar operation will occur for the heart pumps 1 and100B.

In this example, at step 400, the controller 190 monitors axial positionof the impeller 130A using the sensor 195, and determines if theimpeller has moved at step 410, as a result of a change in pressuredifferential between the pulmonary and systemic circulatory systems.This typically involves monitoring signals from the sensor 195 todetermine if a separation between the impeller 130A and the sensor 195has altered, thereby signifying that the impeller 130A has moved from acurrent balance position. The current balance position may correspond tothe normal balance position if the circulatory systems were previouslyfunctioning in accordance with normal haemodynamics, although this isnot essential.

At step 420, the controller determines a direction of the axialmovement, by determining for example if the separation between thesensor 195 and the impeller has increased or decreased, and uses this tocause the magnetic bearing to cause opposing movement of the impeller130A, at step 430, thereby moving the impeller back 130A towards thebalance position.

At step 440, the controller determines the indicator based on rate ofchange of current used by magnetic bearing 180 as the impeller 130A ismoved back towards the balance position. The indicator is compared to apredetermined value at step 450, to determine if the indicator, andhence the power consumption by the magnetic bearing 180 is acceptable atstep 460. The value will therefore typically represent a minimal powerusage by the magnetic bearing 180 that satisfies zero powerrequirements, and this is typically previously determined and stored inthe memory 301 of the controller 190, prior to initial operation of theheart pump 100.

It will be appreciated that in general the power consumed by themagnetic bearing 180 decreases as the impeller 130A nears a new balanceposition associated with the changed pressure differential between thecirculatory systems, and accordingly, this can be used to indicate ifthe impeller 130A has reached a new balance position.

If it is determined that the indicator is not successful at step 460,then the process moves on to step 470 to determine the axial position ofthe impeller 130A. The axial position is then compared to positionlimits at step 480 to determine if the impeller 130A is withinoperational positional limits at step 490. This is performed to ensurethat the impeller 130A does not become too close to the first or secondcavity surfaces 123, 124A, which could impact on the impellerperformance. Again, the positional limits are typically previouslydetermined values stored in the memory 301.

In the event that the impeller 130A is still within positional limitsbut has not yet reached the new balance point, then the process returnsto step 430, to continue causing movement of the impeller 130A.

Otherwise, in the event that it is determined either that the newbalance point has been reached, at step 460, or if positional limitshave been reached at step 490, then movement of the impeller 130A ishalted at the current position, which therefore represents an axialposition as close to the balance point in a zero power configuration asis achievable.

The process can then return to step 400 allowing the process to berepeated in the event that further relative pressure changes occurbetween the systemic and pulmonary systems.

As described above, the configuration is such that the new balance pointwill counteract any variation away from a normal balance of left/rightflow, thereby returning the circulatory systems to the normal flowbalance required by normal haemodynamics.

If applied to the heart pump 100B, the above described processcounteracts any pressure changes within the circulatory systems,returning the system to the normal haemodynamic pressure.

An example of the drive 170 and the magnetic bearing 180 for the heartpumps 100A, 100B is shown in FIGS. 10A to 10F.

In this example, the magnetic bearing 180 includes three evenly spaced120° U-shaped bearing stator cores 181. Turns of copper wire 182 arewound about a radially outer foot of each stator core 181 to produceflux directed towards an iron core 185 attached to the impeller 130 (notshown in this example). A custom NdFeB permanent magnet 183 may beattached to the radially inner foot of each core to provide bias flux tothe magnetic flux path. This permanent magnet may also be located on aimpeller 130, or in the section bridging the inner and outer pole feet184, and should preferably be of high magnetic strength (grade N52). Theaxial bearing magnetically couples to the iron core 185, which is madeof a ferromagnetic material, and is mounted in the impeller 130.

The drive 170 typically includes a slotted axial flux motor coreconstructed with up with 12 poles, but is preferably constructed from astator 171 having six poles 173 on which concentrated copper coils 172are wound. The motor stator 171 is electromagnetically coupled tocircumferentially spaced permanent magnets 174 having opposingpolarities that alternate between North 174 a and South 174 b Poles, andare attached to a ferromagnetic core 175, mounted in the impeller 130.

Each of the bearing stator 181, motor stator 171 and ferromagnetic cores175, 185 are preferably composed of material that exhibits highelectrical resistivity and high magnetic permeability, such as aniron-cobolt or iron-silicon alloy. A suggested material is VACOFLUX 48(Vacuumschmelze GMBH & CO. KG, Germany). The material may be laminatedto reduce eddy current losses.

Axial attractive force f_(z), and the motor torque τ_(z) produced by themagnetic drive and bearing systems 170, 180 are derived from Eq. (1) andEq. (2). The parameters used in each equation are listed in TABLE 1.

TABLE 1 Parameters of the equations B_(R) The peak flux density of themotor B_(S) The peak flux density of the stator M The pole pair numberof the motor ω The rotating speed of the impeller φ The phase differencer₁ The inner radius of the impeller r₂ The outer radius of the impellerz The equivalent distance of the air gap μ₀ The permeability of vacuum

$\begin{matrix}{f_{z} = {\frac{\left( {r_{2\;}^{2} - r_{1}^{1}} \right)\pi}{4\mu_{0}}\left\lbrack {B_{R}^{2} + {2B_{R}B_{S}\cos \; \varphi} + B_{S}^{2}} \right\rbrack}} & (1) \\{\tau_{z} = {\frac{{{zM}\left( {r_{2}^{2} - r_{1}^{2}} \right)}\pi}{2\mu_{0}}B_{R}B_{S}\sin \; M\; \varphi}} & (2)\end{matrix}$

The stator (B_(s)) and the permanent magnet flux of the impeller (B_(r))are assumed to follow a cosine waveform magnetic flux density and asdescribed by Eq. (3) and Eq. (4)

B _(S)(θ,t)=B _(S) cos(ωt−Mθ)  (3)

B _(r)(θ,t)=B _(R) cos(ωt−Mθ−φ)  (4)

The number of turns in the coils 172, 182 and the geometric parametersof the permanent magnets 174, 183 produce this flux.

The operation principle of axial position control by the magneticbearing is shown in FIG. 11A. Changing the magnitude of the motor statorand the magnetic bearing electromagnetic flux B_(S) 176, 186 can controlthe axial displacement of the impeller 130. The motor torque is alsocontrolled changing the phase difference φ.

The permanent magnets 174, 183 in the drive 170 and the bearing 180,each produce a static bias flux 177, 187 in order to reduce the powerrequirements of the system. The attractive force produced by thesemagnets 174, 183 is in balance when the impeller 130 is located in themiddle of the cavity 120. Therefore, the control flux 176, 186 producedby the coils 172, 182 in the drive 170 and the bearing 180,respectively, is required only to stabilise the impeller's axialposition and overcome disturbance forces.

The control flux 186 generated by the bearing 180 can increase ordecrease the effective attraction between the impeller 130 and thebearing 180 by generating a field that is complementary to or counter tothe magnetic field generated by the permanent magnet 183. This controlsthe net magnetic field between the bearing and the bearing magneticmaterial, which in turn allows the position of the impeller 130 to becontrolled in either direction within the cavity 120.

Feedback control of the impeller's axial position is achieved inresponse to impeller displacement, which may be detected by threepositional sensors 195A, 195B, 195C, such as eddy current sensors (U5B,Lion Precision, Minn., USA). These displacement measurements arefeedback parameters used by a control algorithm to stabilize the system.Control gains are output to a power amplifier, which generates therequired current in the corresponding coil to alter the flux density inthe magnetic gap and thus attractive force to maintain impellerlevitation. Similarly, the motor controller generates the three phasecurrent for the motor coils to provide synchronous rotation. Thisrotation may use feedback parameters of rotational speed derived fromhall effect sensors or back EMF recordings to control rotational speed.

The operation principle of tilt control by the magnetic bearing is shownin FIG. 11B. The constant bias flux 177, 187 is produced by thepermanent magnets 174, 183. When the impeller tilts, the electromagneticcontrol flux 176, 186 produced by the axial bearing 180 in the side thatthe impeller 130 approaches is decreased, by applying a counter flux 186_(counter), while the flux in the other side is increased, by applying acomplementary flux 186 _(comp), thereby generating a restoring tiltforce F_(tilt).

It will be appreciated that similar drive and bearing arrangements canbe implemented in the heart pump 1 of FIGS. 1 to 5, albeit with thesebeing arranged between the cavity portions 2, 3, as previouslydescribed.

The manner in which the impeller position is maintained will now bedescribed with reference to FIGS. 12A to 12C. This will be describedwith reference to second and third example heart pumps 100A, 100B, ofFIGS. 6A and 6B. Accordingly, the following description will focusgenerally on a heart pump 100, having a impeller 130, with therespective heart pump 100A, 100B being identified only when this has animpact on the process. It will also be assumed that the heart pump 1functions in a manner substantially similar to that described withrespect to the heart pump 100A, with the impeller 5 acting as theimpeller 130, and this will not therefore be described in furtherdetail.

The controller 190 attempts to maintain a stable axial position of theimpeller 130 in the centre of the pump cavity 120, as shown in FIG. 12A.In this regard, with the impeller centrally positioned, the static biasflux 177A, 187A produced by these magnets 174, 183 is in balance andconsequently only minimal control flux 176A, 186A is required to beproduced by the drive 170 and bearing respectively.

The introduction of a static disturbance force F, shown in FIG. 12B,attempts to displace the impeller 130 from this central position, inthis case towards the cavity portion 121. A slight deviation of theaxial position of the impeller 130 engages the ‘zero position’controller 190, which increases the magnetic control flux 186B generatedby the bearing 180, in order to counteract these forces F and maintainthe notional centralised ‘zero’ position resulting in an increase inmagnetic power used by the magnetic bearing 180.

However, by implementing a ‘zero power’ controller 190, the notionalcentralised axial position of the impeller 130 is not fixed to aphysically central position within the cavity 130, and will change inresponse to the external disturbance forces.

In the above example, the application of a disturbance F toward thecavity portion 121 will cause the controller 190 to move the impeller130 in the opposite direction, toward the bearing 180, until thisdisturbance force is counteracted by the increase in permanent magneticattraction of the bearing bias flux 187C and subsequent reduction ofmotor bias flux 177C. This allows the electromagnetic control flux 186Bto be reduced until the motor and bearing control flux 176C, 186C areminimal and are once again used only to stabilise the impeller in thenew axial position.

As briefly mentioned above, controlling the magnetic bearing in zeropower mode may be advantageous when operating the heart pump in thecardiovascular system, which is constantly adapting and changing itsphysiological parameters. These parameters particularly affect thepressure development within the heart pump, which impose a staticdisturbance hydraulic force on the impeller in the axial and radialdirections.

The ability to alter the hydraulic output of a VAD or BiVAD is importantto enable effective physiological operation within the cardiovascularenvironment. Changes in perfusion requirements due to alterations in apersons physical state must be met by the heart pump. Furthermore, thepossibility of heart chamber collapse due to over-pumping of the heartpump must be prevented, or rectified shortly after such an event occurs,by reducing the heart pump outflow.

The most common technique to achieve this alteration of hydraulicperformance is a change in impeller rotational speed. However, axialdisplacement of the impeller within the pump cavity also changes thehydraulic output of the pump, by inducing changes in flow leakage fromthe high pressure outlet to low pressure inlet. This techniqueeffectively alters the energetic efficiency of the impeller vane set,and is most effective when a semi-open (i.e. without top shroud)impeller is used.

With reference to a single VAD, such as the heart pump 100B depicted inFIGS. 13A to 13C, centrally locating an impeller 54 within the pumpcavity 50 and operating at a set rotational speed about an axis 59, thedesired haemodynamics are produced for the circulatory system in need ofsupport, as shown in FIG. 13B. Without changing the speed, shifting theimpeller along its rotational axis 59 away from inlet 57 as shown by thearrow 51 c effectively increases the clearance above the impeller vanes.This motion reduces the VAD outflow at the outlet 58 as shown by thearrow 53 c, in FIG. 13C. The reduction of outflow is a direct result ofan increase in leakage from the high pressure outlet 58 to low pressureinlet 57, as shown by the arrows 55 c. Similarly, shifting the impellertoward the inlet 57, as shown by the arrow 51 b has the opposite effect,i.e. VAD outflow increases, as shown by the arrow 53 b, in FIG. 13A.

The axial motion of the impeller 130B can therefore be used to deliversufficiently variable output flow from the VAD to meet the physiologicalrequirements of the cardiovascular system at a set rotational speed.This effect is more pronounced when the impeller blade height is low andthe ratio of impeller blade height to axial clearance is approx 3:1(blade height=1.4 mm, starting axial clearance gap=0.5 mm), andsufficient outflow variation achieved with a movement of +/−0.3 mm.

In a specific application, the ability to acutely alter the left andright outflow of a BiVAD system is important, especially in the postoperative period, to accommodate relative changes in systemic andpulmonary vascular resistance, the relative level of ventricularcontractility, alleviation of acute pulmonic or systemic congestion, andto prevent suckdown by maintaining adequate atrial filling pressures.This requirement is again achieved by axially displacing the impeller130A within a dual chambered heart pump, such as the heart pump 100A,thereby maintaining a long term balance of left and right flows.

With reference to FIGS. 14A to 14C, an example of the effect of axialposition on relative flows for a BiVAD, such as the heart pumps 1, 100Awill now be described. This example is described with respect to theheart pump of FIG. 6A, and accordingly, similar reference numerals areused to designate similar features, although it will be appreciated thatthe techniques are equally applicable to the heart pump 1 of FIGS. 1 to5.

In this example, with the impeller 130A located within the pump cavity120, at the physical axial centre of the cavity, as shown by the line20, and operating at a set rotational speed of approximately 2300 rpmabout an axis 19, this results in the desired haemodynamics of 100 mmHg(LVAD) and 20 mmHg (RVAD) being produced for the systemic and pulmonarysystems. Accordingly, the flows via the first and second outlets 151,152 are in balance at approximately 5 L/min, as shown by the arrows 23a, 24 a in FIG. 14B. The exact rate of flow from the left cavity isslightly higher than the right cavity, due to the natural outflowdifferential of the heart caused by the bronchial circulation.

Without changing the speed, shifting the impeller 0.3 mm along itsrotational axis 19 toward the second cavity portion 122A, as shown bythe arrow 21 c reduces the axial clearance above the RVAD vanes (to 0.2mm), while simultaneously increasing the clearance above the LVAD vanes(0.8 mm). This motion improves the RVAD outflow via the outlet 152, asshown by the arrow 24 c while reducing the LVAD outflow via the outlet151, as shown by the arrow 23 c, in FIG. 14C. The reduction of outflowis a direct result of an increase in leakage from the high pressureoutlet 151 to low pressure inlet 141, as shown by the arrows 25 c.

Similarly, shifting the impeller from the central position toward thefirst cavity 121, as shown by the arrow 21 b has the opposite effect.For example, while maintaining arterial pressures, this 0.3 mm movementtoward the LVAD cavity 121 increases left outflow to 6.4 L/min, as shownby the arrow 23 b while right outflow reduces to 4.6 L/min, as shown bythe arrow 24 b, in FIG. 14A, representing an instantaneous flowdifferential of 1.8 L/min (36%). This axial motion can therefore be usedto accommodate the required variable flow output from left and righthearts at a set rotational speed. This effect is more pronounced whenthe impeller blade height is low and the ratio of impeller blade heightto axial clearance is approx 3:1.

A graph showing examples of the relative performances of the pump, andin particular the pressure and flow at the outlets 151, 152, is shown at1500, and 1510 respectively, in FIG. 15.

An example BiVAD configuration heart pump similar to the heart pump 100Aof FIG. 6A, will now be described, with additional features being shownin FIGS. 16A to 16H. It will be appreciated that similar features mayalso be implemented in VADs, where appropriate.

In this example, the heart pump draws inflow from the right atrium viathe second inlet 142 and from the left ventricle via the first inlet141, and provides outflow to the pulmonary artery via the second outlet152 and to the aorta via the first outlet 141. In this example, thefirst and second sets of vanes 131, 132 are mounted on a shared rotatinghub to form a magnetically and hydro-dynamically suspended centrifugalimpeller 130A. The first and second sets of vanes 131, 132 have adifferent outer diameter to produce the pressure required of thesystemic and pulmonary systems at a common rotational speed, asdescribed above. The differential in flow required from the left andright hearts is achieved by alteration of axial clearance above eachsemi-open VAD impeller.

The suspension system incorporates a hydrodynamic journal bearing 115,whilst volutes 111, 112 are provided as part of the housing 110, tothereby assist with transfer of the fluid to the outlets 151, 152. Thevolutes maybe any combination of type spiral/single, split/double orcircular/concentric, however the latter circular volute type ispreferred, as this configuration produces a stabilising radial hydraulicforce for optimal journal bearing functionality.

It will be appreciated that in the heart pump 1 of FIGS. 1 to 5 ahydrodynamic bearing could be implemented as part of the shaft 6. In oneexample, to achieve this, a separation between the shaft 6 and theconnecting tube 4 is configured to be between 0.05 mm and 0.1 mm.

As outlined above, axial hydraulic force is imposed on the impeller 130Adue to the build up of pressure within the left and right cavities. Adifferential of these pressures acting on the top and bottom impellerfaces will produce a resultant force in either the positive or negativeaxial direction.

For the condition of dual heat support, the most important parameter tomaintain is the balance of left and right flows from each supportingpump. Mismatch of flows may lead to the potentially disastrous situationof left or right heart chamber collapse. Application of the zero powercontroller, and its ability to favourably adapt device performance isdescribed below for a number of foreseeable physiological conditions tobe encountered by the BiVAD.

FIGS. 17A to 17D shows the resultant axial force development on a doublesided impeller 130A during a variety of common conditions. Whilst thisis described with respect to the impeller 130A, the techniques areequally applicable to the impeller 5 of FIGS. 1 to 5.

During diastole, the force acting on the face of the semi open LVADimpeller decreases from the outer diameter to the inner diameter, inproportion to the pressure development along this path. This is balancedby the force acting beneath the impeller 130A, which is composed of thesection exposed to the high LVAD outlet pressure (which does not reduceas much with reducing diameter), and the RVAD impeller section exposedto the lower pressure developed in the RVAD cavity.

This balance is destroyed during systole, as shown in FIG. 17B. In thisexample, the left ventricular pressure acts across the entire LVADimpeller face, which cannot be matched by the lower RVAD pressure. Dueto the impulse nature of this disturbance and damping of the blood, theactual force to be countered by the magnetic bearing is lower than thestatic differential would suggest.

The effect of an increase in systemic vascular resistance (SVR) onpressure development and thus force generation is described in FIGS.17A, 17B and 17C. Increasing SVR acts to increase the pressure withinthe LVAD first cavity 121, which in turn produces a force toward theRVAD second cavity 122A. The same resultant force is produced for adecrease in pulmonary vascular resistance (PVR). Using this logic, anopposite force is produced toward the LVAD cavity when SVR is decreasedand/or PVR is increased.

Finally, the force generated on the impeller during instances of leftheart chamber collapse is described in FIG. 17D. In this instance, thepressure generated in the inlet cannula drops well below 0 mmHg, and theoutlet correspondingly drops to a low value of pressure. Thiscombination contributes to a large axial force toward the LVAD cavity.

FIGS. 18A to 18C describe the adaptation of the heart pump 100A toalterations in relative vascular resistance. The example presented isfor a relative increase in SVR compared to PVR, but could also depict arelative decrease in PVR compared to SVR. The exact opposite would occurwith a relative decrease in SVR or increase in PVR. Furthermore, whilstthis is described with respect to the heart pump 100A, the techniquesare equally applicable to the heart pump 1 of FIGS. 1 to 5.

During normal operation shown in FIG. 18A, the impeller 130A ispositioned centrally within the cavity 120, and the left and rightoutflows from the first and second inlets 151, 152 are in balance, asshown by the arrows 36 a, 35 a. As the impeller 130A is located in thecentre of the pump cavity 120, the bearing and motor PM bias forces,shown by the arrows 34 a, 31 a, are also in balance. Electromagneticcontrol flux forces from the drive 170 and the bearing 180 are minimal,as shown by the arrows 32 a, 33 a, being required only for stabilisationand to account for dynamic disturbances.

However, an increase in SVR, as shown in FIG. 18B, causes a decrease inLVAD flow from the outlet 151, as shown by the arrow 36 b, and anincrease in LVAD cavity pressure, which, as described earlier in FIG.17C, results in a static axial hydraulic force vector toward the RVADcavity 122A. To maintain the impeller 130A in the central position, themagnetic bearing flux must increase in magnitude (and thus power) togenerate a counter field opposing that of the permanent magnets 183, asshown by the arrow 33 b, to provide a restoring force.

However, this elevated SVR must be overcome to restore LVAD outflow.Implementing the ‘Zero power’ controller as described above, will causethe impeller to move toward the LVAD cavity 121 until the disturbanceforce shown by arrow 30 is balanced by the increase in permanent magnetbias flux from the drive 170, as shown by the arrow 31 c. In thisequalised position, electromagnetic control flux is once again minimal,and required only for dynamic disturbance forces, thus power consumptionis reduced. Most importantly however, LVAD flow from the outlet 151increases, as shown by the arrow 36 c, whilst RVAD flow from the outlet152, as shown by the arrow 35 c slightly decreases, and the balance ismaintained. Rotational speed may also be increased to simultaneouslyincrease both LVAD and RVAD flow, should this be required to maintainabsolute flow levels.

FIGS. 19A to 19C describe the adaptation of the heart pump 100A to heartchamber collapse. The example presented describes the consequences ofleft heart chamber collapse, however the opposite characteristics wouldoccur for right heart chamber collapse. Furthermore, whilst this isdescribed with respect to the heart pump 100A, the techniques areequally applicable to the heart pump 1 of FIGS. 1 to 5.

In the example of FIG. 19A, left heart chamber collapse impairs the flowof blood into the inlet 141 on left side of the heart pump, as shown at47, resulting in a severe reduction of LVAD outflow from the outlet 151,as shown by the arrow 46 a. As shown in FIG. 17D, an axial force vectortoward the LVAD cavity ensues, which must be counteracted by an increasein bearing magnetic control flux, as shown by the arrow 43 a, to therebymaintain the centralised impeller position.

However, the ‘Zero Power’ control method further automatically adjuststhe impeller axial position toward the RVAD cavity 122A, as shown inFIG. 19B, until the bearing permanent magnet bias force shown by arrow44 b balances the disturbance hydraulic force 40. This action returnsthe electromagnetic control flux generated by the drive 170 and thebearing 180, as shown by the arrows 43 b, 42 b to minimal levels, thusreducing power consumption. Furthermore, RVAD outflow from the outlet152 is increased and the pressure differential from LVAD inlet 141 tooutlet 151 is decreased. This subsequently shifts blood into thepulmonary circuit, and consequently to the left heart chamber, thusalleviating the collapse of the left heart. LVAD and RVAD cavitypressures then return to a normal state, thus the disturbance force iseliminated, causing the impeller 130A to automatically translate back tothe centre of the cavity 120, which in turn returns the balance of LVADand RVAD outflows from the outlets 151, 152, as shown by the arrows 46c, 45 c.

In another example, a single VAD heart pump, such as the heart pump100B, is adapted to provide ventricular assistance to one side of afailing heart.

In this example, axial hydraulic force is imposed on the impeller 130Bdue to the build up of pressure within the first and second cavities121, 122B. A differential of these pressures acting on the top andbottom impeller faces will produce a resultant force in either thepositive or negative axial direction. The pressure differential arisesdue to flow between the inlet 141 and outlet 151, which can result in apressure gradient across the face of the impeller 130B that includes thevanes, whereas the pressure gradient on the other face is substantiallyconstant.

FIGS. 20A to 20D show the resultant axial force development on theimpeller during a variety of common conditions. During diastole, asshown in FIG. 20A, the force beneath the impeller 130B is only partiallybalanced by the force acting on the top of the impeller, resulting in anet force towards the inlet 141, shown by the arrow 2000. This forcedecreases from the outer diameter to the inner diameter, in proportionto the pressure development along this path. This imbalance is reducedduring systole, as shown in FIG. 20B, when the left ventricular pressureacts across the entire VAD impeller face.

The effect of an increase in systemic vascular resistance (SVR) onpressure development and thus force generation is described in FIG. 20C.Increasing SVR acts to increase the overall pressure in the pump cavity,which in turn increases the mismatch of force acting beneath and abovethe impeller 130B, resulting in a net force shown by arrow 2010, towardsthe inlet 141. Finally, the force generated on the impeller duringinstances of heart chamber collapse is described in FIG. 20D. In thisinstance, the pressure generated in the inlet cannula drops well below 0mmHg, and the outlet correspondingly drops to a low value of pressure.This combination contributes to an increase in axial force toward theinlet 141, shown by the arrow 2020.

For the single heart support application, the two parameters in mostneed of control are the outflow pressure (and thus flow) in response toalterations in vascular resistance, and the rectification of heartcollapse. The ability for the zero power controller to favourably adaptheart pump performance is described below for a number of foreseeablephysiological conditions to be encountered by a single VAD.

FIGS. 21A to 21C describe the adaptation of the heart pump 100B toincreases in vascular resistance.

In FIG. 21A, during normal operation, the heart pump is configured sothat with the impeller 130B positioned approximately centrally withinthe cavity 120, the outflow pressure and flow provided at the heart pumpoutlet 151 meets the circulation system's physiological requirements, asshown by the arrow 63 a. In this configuration, with the impeller 130Blocated in the centre of the pump cavity 120, the bearing 180 and drive170 PM bias forces shown by the arrows 67 a, 64 a, are in balance.Electromagnetic control flux from the drive 170 and the bearing 180 isminimal, as shown by the arrows 65 a, 66 a, and required only forstabilisation and to account for dynamic disturbances.

An increase in vascular resistance (VR), shown in FIG. 21B, causes anincrease in VAD outlet pressure as shown by the arrow 63 b and thuscavity pressure, which, as described in FIG. 20C, results in an increaseof static axial hydraulic force toward the first cavity 121, as shown bythe arrow 68. To maintain the impeller 130B in the central position, themagnetic bearing flux must increase in magnitude (and thus power), asshown by the arrow 66 b, to thereby provide a restoring force.

However, in the single VAD application, this increase in outlet pressuredue to the elevated VR should be reduced, as per the naturalbaroreceptor reflex. Alterations in flow can be achieved with speedchanges, while flow balancing problems are negated by the ability forthe remaining functioning ventricle's ability to balance the flow.Implementing the ‘Zero power’ controller as described above, will causethe impeller to move away from the first cavity 121 until thedisturbance force 68 is balanced by the increase in PM bias flux fromthe bearing 180, as shown by the arrow 67 c. In this equalised position,electromagnetic control flux is once again minimal, and required onlyfor dynamic disturbance forces, thus power consumption is reduced. Mostimportantly however, VAD outflow pressure shown by the arrow 36 creturns to a lower value, while outflow reduces (due to the increasedvascular resistance). The opposite characteristic is observed for adecrease in vascular resistance (i.e. impeller movement toward theinlet), as encountered in a state of exercise. Thus the automaticimpeller movement maintains vascular pressure, and thus outflowincreases to suit.

However, this increase in flow may lead to a situation of overpumpingand thus heart chamber collapse. FIGS. 22A to 22C describe theadaptation of the heart pump 100B to this event.

When heart chamber collapse occurs, as shown in FIG. 22A, this impairsthe flow of blood into the inlet 141, as shown by the arrow 79,resulting in a severe reduction of VAD outflow at the outlet 151, asshown by the arrow 73 a. As described in FIG. 20D, an axial force vectortoward the inlet 141 is developed, which must be counteracted by anincrease in bearing magnetic control flux, shown by the arrow 76 a, tomaintain the centralised impeller position.

However, the ‘Zero Power’ control method automatically adjusts theimpeller axial position away from the inlet cavity 141, until thebearing PM bias force shown by the arrow 77 b balances the disturbancehydraulic force shown by the arrow 78, in FIG. 22B. This action returnsthe electromagnetic control fluxes shown by the arrows 75 b, 76 b tominimal levels, thus reducing power consumption. The pressuredifferential across the VAD inlet to outlet is therefore decreased, thusalleviating the collapse of the left heart. VAD cavity pressures thenreturn to a normal state, eliminating the disturbance force and causingthe impeller to automatically translate back to the centre of the cavityand returning outflow to normal, as shown by the arrow 79 c in FIG. 22C.

Experiments were performed using a heart pump similar to that shown inFIG. 6A, configured to operate in accordance with the control process ofFIG. 9. The heart pump was coupled to a fluid circulation loop designedto simulate various hemodynamic conditions, such as PulmonaryHypertension, LVAD inflow obstruction/left ventricular suckdown, andSystemic Hypertension, thereby allowing the responsiveness of the pumpto be assessed.

All resulting parameters and haemodynamics are described in table 2, andillustrated in FIGS. 23A to 23E, with a detailed description providedsubsequently. The conditions simulated can be summarised as follows:

-   -   1. Normal Condition (FIG. 12A)    -   2. Pulmonary Hypertension (FIG. 12B)    -   3. Impeller movement RIGHT to account for PHT (FIG. 12C)    -   4. Normal Condition    -   5. Suckdown event (19A)    -   6. Impeller movement RIGHT to account for Suckdown (19B)    -   7. Normal Condition (18A)    -   8. Systemic Hypertension (18B)    -   9. Impeller movement LEFT to account for SHT (18C)

TABLE 2 Condition 1 2 3 4 5 6 7 8 9 Z* (mm) 0 0 −0.2 0 −0.1 −0.6 0 −0.10.5 AoP (mmHg) 98 100 96 98 83 75 98 150 180 PAP (mmHg) 17 20 22 17 2630 17 20 16 LAP (mmHg) 9 7 10 9 3 9 9 14.5 8 RAP (mmHg) 6 7 5 6 10 5.5 61.5 5 SQ (L/min) 5.1 5 5 5.1 4.6 4.4 5.1 4.2 4.5 PQ (L/min) 5.6 5.6 5.35.6 4.7 4.4 5.6 4.9 5.5 SVR 1600 1620 1565 1600 1450 1340 1600 2950 3320(dynes · s · cm⁵) PVR 94 168 163 94 370 340 94 78 91 (dynes · s · cm⁵)Hydraulic 4.2 5 4.2 4.2 12 10.1 4.2 2.7 2.5 Force* (N) Bearing Power 00.672 0.2625 0 20 2.3625 0 2.3625 0.2625 (Watts) *Indicates a positivedirection toward the LVAD cavity, Z = impeller axial position (mm), AoP= Aortic Pressure, PAP = Pulmonary Arterial Pressure, LAP = Left AtrialPressure, RAP = Right Atrial Pressure, SQ = Systemic Flow Rate, PQ =Pulmonary Flow Rate, SVR = Systemic Vascular Resistance, PVR = PulmonaryVascular Resistance.

In the first example, pulmonary hypertension was simulated. In practice,such an event occurring in the human body would result in a reduction offlow through the pulmonary system, thus reducing venous return to theleft heart and a consequential increased potential for suck-down of theleft heart chambers. An example of the response of the heart pump insuch situations is as described above with respect to FIGS. 12A to 12C.

Referring to FIG. 12A, in an initial “normal” condition (Condition 1)the impeller 130A is positioned centrally with normal pressures andflows being produced at a rotational speed of 2500 rpm. This is achievedusing an impeller 130A having a diameter of 50 mm in the first LVADcavity 121 and approximately 25 mm in the second RVAD cavity 122A, withthe vanes having a height of 1.4 mm, and a starting axial clearance gapof 0.5 mm, to thereby allow axial impeller movement of +/−0.5 mm.

In this case, the hydraulic force caused by the differential pressurefrom the RVAD and LVAD cavities 122, 121, respectively, equals +4.2Ntoward the LVAD cavity 121. This is immediately balanced by the magneticbias force 187A caused by the permanent magnets within the magneticbearing 180. Bearing current is only required for disturbance forces andthus the bearing electromagnetic force 186A and thus power usage equalszero, as shown in Table 2.

An incidence of Pulmonary Hypertension is then simulated (Condition 2),which is as described in FIG. 12B with results being illustrated inFIGS. 23A to 23E. The haemodynamic result is an immediate decrease inflow through the pulmonary system and left atrial pressure from 9 to 7mmHg. In this case, the hydraulic forces F toward the LVAD/Motor areincreased by 0.8N to +5.0N, due to the elevated pulmonary vascularresistance (168 dynes.s.cm⁵). To maintain the impeller 130A in thecentral position, the controller 190 inputs current into the magneticbearing 180 to produce an additional electromagnetic force 186B of 0.8N,thus balancing all external forces at the expense of an additional 0.672W bearing power. If this condition remains, the reduced flow through thepulmonary system may further reduce left atrial pressure to inducesuckdown in the left heart.

To avert this situation, the controller 190 then creates an “automaticresponse” condition (Condition 3) and FIG. 12C, whereby the impeller130A is automatically moved toward the RVAD cavity −0.2 mm. In doing so,pulmonary arterial pressure increases as the RVAD improves its hydraulicefficiency, pushing more flow to the left heart to halt the reduction inleft atrial pressure. This reduces the hydraulic force F on the impeller130A by 0.8N, back to +4.2N. Motion of the impeller 130A towards themagnetic bearing 180 causes the magnetic bias force 187C created by thepermanent magnets within the magnetic bearing 180 to increase inmagnitude, to 5.3N. A small cancellation current of −0.25 amps istherefore required to pass through the bearing coils to reduce thebearing force to +4.2N and restore the force balance. This results in abearing power of 0.26 W, which is less than the bearing power prior tomovement (0.672 W) thus demonstrating the effect of the controller 190to minimise bearing power when operating in a “zero power” mode asdescribed above. These results highlight that the hemodynamic responseof the pump functions as expected, and that the heart pump uses minimumpower when in the automatic response state (Condition 3) required tocounteract the pulmonary hypertension state.

In the second example, LVAD inflow obstruction was simulated. Inpractice, such an event occurring in the human body would result in areduction of flow through the pulmonary system (due to increased PVR),and a consequential increased potential for suck-down of the left heartchambers. An example of the response of the heart pump in suchsituations is as described above with respect to FIGS. 19A to 19C.

The heart pump 100A and circulatory system are in an initial “normal”condition (Condition 4), which is substantially as for the normalcondition (Condition 1) described above, and will not therefore bedescribed in any further detail.

An incidence of left ventricular suckdown/obstruction is then simulated(Condition 5), as described in FIG. 19A. The haemodynamic result is animmediate decrease in left atrial pressure. In this case, the hydraulicforces 40 toward the LVAD/Motor are increased. To maintain the impellerin the central position, the controller inputs current into the magneticbearing to produce an additional electromagnetic force (43 a).

The controller 190 then seeks a minimal power position, which takes theimpeller 130A in the direction of the RVAD cavity 122 by −0.1 mm. Inthis position, an electromagnetic force 43 a of 12N is still required tobalance all external forces, at the expense of 20 W of bearing power. Asa result, the LAP reduces from 9 mmHg to 3 mmHg. If this conditionremained, the reduced pulmonary flow may further reduce left atrialpressure to induce complete suckdown in the left heart.

The controller 190 then implements the automated response situationshown in FIG. 19B (Condition 6), by moving toward the RVAD cavity 122 tothe position −0.5 mm, thereby minimising the power used by the magneticbearing. In doing so, both aortic pressure and flow decrease as the LVADreduces its hydraulic efficiency, whilst both pulmonary arterialpressure and flow increase as the RVAD improves its hydraulicefficiency, pushing more flow to the left heart to halt the reduction inleft atrial pressure and raise it back to 9 mmHg and thus rectifying thesuckdown event.

This motion decreases the force on the rotor by 1.9N (40) to +10.1N. Theimpeller motion toward the magnetic bearing causes the magnetic biasforce 44 b created by the permanent magnets within the magnetic bearingto increase in magnitude. A small cancellation current of −0.75 ampsonly is now required to pass through the bearing coils to reduce thebearing force to +10.1N and restore the force balance. This results in abearing power of 2.36 W, which is less than the bearing power prior tomovement (20 W) thus demonstrating the effect of the zero powercontroller to minimise bearing power.

A further slight motion back toward the LVAD cavity 121 would then finetune the impellers operating position by increasing the permanentmagnetic bias force and return a true zero power reading.

In a third example, systemic hypertension was simulated. In practice,such an event occurring in the human body would result in a reduction offlow through the systemic system, thus reducing venous return to theright heart and a consequential increased potential for suck-down of theright heart chambers. An example of this is shown in FIGS. 18A to 18C.

In FIG. 18A, the heart pump 100A and circulatory system are in aninitial “normal” condition (Condition 7), which is substantially as forthe normal condition (Condition 1) described above, and will nottherefore be described in any further detail.

An incidence of Systemic Hypertension is then simulated (Condition 8),as shown in FIG. 18B. The haemodynamic result is an immediate decreasein right atrial pressure. In this case, the hydraulic forces toward theLVAD cavity 122 are decreased. Thus more electromagnetic power would berequired by the magnetic bearing 180 to increase the magnetic bearingforce 33 b to maintain a set central impeller position.

In doing this, the impeller 130A may be allowed to move −0.1 mm towardthe RVAD cavity 122. In this case, the hydraulic force toward the LVADdecreases by a further 1.5N to +2.7N, due to the excessively elevatedsystemic vascular resistance (2950 dynes.s.cm⁵). To maintain theimpeller 130A in this position, the controller inputs cancellationcurrent into the magnetic bearing to reduce the bias electromagneticforce 34 b by the required 1.5N, thus balancing all external forces atthe expense of an additional 2.36 W bearing power. If this conditionremained, the reduced LVAD outflow may further reduce right atrialpressure to induce suckdown in the right heart.

To avert this situation, the controller 190 then implements theautomated response (Condition 9) as shown in FIG. 18C, whereby theimpeller 130A is automatically moved toward the LVAD cavity 121 by +0.5mm. In doing so, both aortic pressure and flow increase as the LVADimproves its hydraulic efficiency, pushing more flow to the right heartto halt the reduction in right atrial pressure. This decreases thehydraulic force on the rotor by an additional 0.2N (30), to +2.5N. Theimpeller motion toward the drive 170 causes the magnetic bias force 34 ccreated by the permanent magnets within the magnetic bearing to decreasein magnitude. In this case, the total hydraulic force is almost balancedby the permanent magnetic bias force, needing only 0.26 W of magneticbearing power to balance the force. Thus the magnetic bearing current 33c is essentially returned to a minimum, and thus a minimal powercondition is observed.

The above described results highlight that the control process cantherefore automatically adjust heart pump outflow in response to theconditions of the system in which it operates. In particular, the systemuses an axial magnetic bearing and drive to suspend and rotate theimpeller of the centrifugal blood pump, which is operated under a zeropower control condition. This control process acts to automaticallyadjust the axial position of the impeller in response to changing axialhydraulic forces imposed on the impeller. This technique allows theheart pump to mimic the flow balancing property of the Frank-Starlinglaw of the heart and thus automatically adapt to changes in atrialpressure (preload) and vascular resistance (afterload). This isparticularly advantageous for preventing the potentially disastrouscollapse of left or right heart chambers, as well as overcoming changesin vascular resistance.

One example application of the controller relates to its use in aBi-ventricular assist heart pump. This heart pump includes an impellerhaving left and right vanes positioned on a shared rotating hub that iscompletely suspended in the blood. This suspension system incorporatesan electromagnetic motor and axial magnetic bearing system for axialsuspension and drive, while radial support is achieved using ahydrodynamic journal bearing. This journal bearing is well washed by theinherent shunt flow from left to right cavities. The left and rightvanes have a different outer diameter to produce the pressure requiredof the systemic and pulmonary systems at a common rotational speed.

The instantaneous differential in flow required to balance outflow fromthe left and right hearts is achieved by the alteration of axialclearance above these semi-open vanes. Thus, an axial motion toward theleft cavity will reduce the axial clearance gap above the left vane set,thus increasing the left heart outflow. Simultaneously, the gap abovethe right vane set will increase, thus reducing the right heart outflow.Similarly, a motion to the right cavity will induce the opposite effect.

This allows the heart pump to automatically adjust the outflow of theleft and right cavities to minimise the potential for the left or rightheart chambers to collapse, as well as to account for relativealterations in vascular resistance. For example, in the event of heartchamber collapse, the hub will translate away from the collapsed side,thus reducing the suction at the inlet whilst also increasing the flowfrom the opposite pump. When encountering an increase in relativevascular resistance, the impeller will translate toward the side withthe increased afterload, thus enabling the affected side to overcome theresistance and maintain flow balance.

This enables a Frank Starling-like control of this flow balance, whichis achieved automatically by the incorporation of the zero powermagnetic control algorithm.

In another example application, the controller 190 is used for a singleventricular assist heart pump. This heart pump includes an impellerhaving a single set of vanes positioned on a rotating hub that iscompletely suspended in the blood. This suspension system againincorporates an electromagnetic motor and axial magnetic bearing systemfor axial suspension and drive, while radial support is achieved using ahydrodynamic journal bearing. This journal bearing is well washed byvirtue of a hole in the impeller, which allows blood flow along anunderside of the impeller.

Pressure control is achieved by the alteration of axial clearance abovethese semi-open vanes. Thus, an axial motion toward the inlet willreduce the axial clearance gap above the vanes, thus increasing thepressure developed in the pump.

This allows the heart pump to automatically adjust the pressuredeveloped by the pump to minimise the potential of heart chambercollapse, as well as to account for pressure changes within thecirculatory system. For example, in the event of heart chamber collapse,the hub will translate away from the inlet, thus reducing the suction atthe inlet. When encountering an increase in vascular pressure, theimpeller will translate away from the inlet, thus providing pressurerelief.

Accordingly, the above described system can provide a VAD that canautomatically control the outflow in response to the needs of thecirculation system. This uses an axial magnetic bearing system thatimplements a zero power controller that adjusts the axial position ofthe impeller in response to hydraulic force. In one example, the motorand bearing arrangements each result in a net attractive force on theimpeller 130, allowing the impeller to be provided at a balance pointwhose position is dependent on relative pressures in the pump cavity. Bysuitable arrangement, this can be used to provide relative flow controlin BiVAD applications, and relative pressure control in VADapplications.

In one example, a combination of the axial movement (+/−0.3 mm) andsmall impeller blade heights (1-2 mm) produces sufficient change tooutflow hydraulics at an unchanged rotational speed.

This can provide sensitivity to atrial (preload) pressure and arterial(afterload) pressure, as well as allowing rectification of heart chambercollapse. This allows the heart pump to maintain a suitable left/rightflow balance, dependant on atrial pressure, in a BiVAD embodiment, or aset arterial pressure in a left or right VAD embodiment, similar to thebaroreceptor reflex.

Accordingly, this avoids the need to change centrifugal pump rotationalspeed to produce changes in pump performance that is used in traditionalheart pumps. This change in performance is advantageous to meet thephysiological requirements of the circulatory system, whilst avoiding acomplex and active physiological control algorithm that receivesfeedback from hardware sensors and uses software estimation to controlpump speed. These hardware sensors induce further reliability issuesthat limit the long term durability of the heart pump, whilst thesoftware estimation introduces complexity.

The control process also addresses the issue of flow balancing from theleft and right pumps in a bi-ventricular assist system, particularlyfrom a single rotary impeller system.

Accordingly, the control process provides a controller for a rotary typeheart pump that can automatically and passively adjust the hydraulicoutput of its rotating impeller without changing rotational speed, andthat does not rely on feedback from haemodynamic sensors or softwareestimation. Instead, the impeller alters its output in response tochanges in preload and after load, similar to the Frank-Starling law ofthe heart.

In one example, the control process is achieved by the incorporation ofan axial magnetic motor and bearing, which implements a zero powercontroller to automatically adjust the axial position of the centrifugalimpeller within the pump cavity. The zero power controller responds toalterations in axial hydraulic force encountered when pump preload andafterload change. A change in performance with axial movement is mosteffectively observed when the impeller incorporates a set of semi open(unshrouded) blades.

The control process also provides the ability for a Bi-ventricularassist system to automatically adjust its flow rate in order to maintaina suitable left/right flow balance. This is most efficiently achieved ina solitary rotary type centrifugal heart pump implementing the ‘zeropower’ controller and adapted to provide bi-ventricular assistance.

Persons skilled in the art will appreciate that numerous variations andmodifications will become apparent. All such variations andmodifications which become apparent to persons skilled in the art,should be considered to fall within the spirit and scope that theinvention broadly appearing before described.

For example, functionality provided by separate motor and bearingarrangements could be achieved using a combined arrangement, in whichone end of the housing includes a set of passive attractive magnets,whilst the other end of the housing includes a combined motor andbearing windings.

1) A heart pump including: a) first and second cavity portions, eachcavity portion including a respective inlet and outlet; b) a connectingtube extending between the first and second cavity portion; c) animpeller including: i) a first set of vanes mounted on a first rotor inthe first cavity portions; ii) a second set of vanes mounted on a secondrotor in the second cavity portion; and, iii) a shaft connecting thefirst and second rotors, the shaft extending through the connectingtube; d) a drive for rotating the impeller; and, e) a magnetic bearingincluding at least one bearing coil for controlling an axial position ofthe impeller, at least one of the drive and magnetic bearing beingprovided at least partially between the first and second cavityportions. 2) A heart pump according to claim 1, wherein the axialposition determines a separation between each set of vanes and arespective cavity surface, the separation being used to control thefluid flows from the inlets to the outlets. 3) A heart pump according toclaim 1 or claim 2, wherein the drive includes: a) a second magneticmaterial provided in the impeller; b) at least one drive coil that inuse generates a magnetic field that cooperates with the second magneticmaterial allowing the impeller to be rotated. 4) A heart pump accordingto claim 3, wherein the second magnetic material includes a number ofcircumferentially spaced permanent magnets mounted in the impeller,adjacent magnets having opposing polarities. 5) A heart pump accordingto claim 3 or claim 4, wherein the second magnetic material is mountedin the second rotor, and wherein the drive is positioned adjacent thesecond cavity, the drive and second magnetic material being configuredto result in an attractive force between the drive and the second rotor.6) A heart pump according to any one of the claims 1 to 5, wherein, inuse, the at least one bearing coil generates a magnetic field thatcooperates with first magnetic material in the impeller, allowing theaxial position of the impeller to be controlled. 7) A heart pumpaccording to claim 6, wherein the first magnetic material is a permanentmagnet. 8) A heart pump according to claim 7, wherein the at least onebearing coil is for generating a magnetic field that is one ofcomplementary to and counter to the first magnetic field generated bythe permanent magnet, thereby controlling the net magnetic field betweenthe bearing and the first magnetic material. 9) A heart pump accordingto any one of the claims 1 to 8, wherein the first magnetic material ismounted in the first rotor and, wherein the magnetic bearing ispositioned adjacent the first cavity, the magnetic bearing and firstmagnetic material being configured to result in an attractive forcebetween the magnetic bearing and the first rotor. 10) A heart pumpaccording to any one of the claims 1 to 9, wherein the impellerincludes: a) a second magnetic material provided on the second rotor forcooperating with the drive to allow rotation of the impeller; and, b) afirst magnetic material provided on the first rotor for cooperating withthe magnetic bearing to allow the axial position of the impeller to becontrolled. 11) A heart pump according to any one of the claims 1 to 10,wherein: a) the magnetic bearing is positioned adjacent the firstcavity, the magnetic bearing and first rotor being configured to resultin a first attractive force between the magnetic bearing and the secondrotor; and, b) the drive is positioned adjacent the second cavity, thedrive and second rotor being configured to result in a second attractiveforce between the drive and the second rotor and wherein the first andsecond attractive forces are approximately balanced when the impeller ispositioned at an approximately axially central position during normalcirculatory conditions. 12) A heart pump according to any one of theclaims 1 to 11, wherein the heart pump includes a controller for: a)determining movement of the impeller in a first axial direction; b)causing the magnetic bearing to move the impeller in a second axialdirection opposite the first axial direction; c) determining anindicator indicative of the power used by the magnetic bearing; and, d)causing the magnetic bearing to control the axial position of theimpeller in accordance with the indicator to thereby control a fluidflow between the inlets and the outlets. 13) A heart pump according toclaim 12, wherein the controller is for: a) comparing the indicator to athreshold; and, b) causing the magnetic bearing to stop movement of theimpeller in the second axial direction depending on the results of thecomparison. 14) A heart pump according to claim 12 or claim 13, whereinthe controller is for minimizing the power used by the magnetic bearing.15) A heart pump according to any one of the claims 12 to 14, whereinthe controller is for: a) comparing an axial position of the impeller toposition limits; and, b) controlling the magnetic bearing to maintainthe axial position of the impeller within the position limits. 16) Aheart pump according to any one of the claims 12 to 15, wherein thecontroller is for: a) determining a pressure change within at least partof a cavity; and, b) controlling the axial position of the impeller inresponse to the pressure change. 17) A heart pump according to claim 16,wherein the controller is for determining the pressure change bydetecting axial movement of the impeller. 18) A heart pump according toany one of the claims 12 to 17, wherein the controller is for: a)detecting movement of the impeller caused by a change in fluid pressurewithin at least one of the cavity portions; and, b) causing the magneticbearing to control the axial position of the impeller to thereby changea fluid flow from the inlet to the outlet for at least one of the cavityportions. 19) A heart pump according to claim 18, wherein the controlleris for, at least one of: a) causing the magnetic bearing to reduce theseparation between the vanes and the cavity surface to thereby increasethe flow of fluid from the inlet to the outlet; and, b) causing themagnetic bearing to increase the separation between the vanes and thecavity surface to thereby decrease the flow of fluid from the inlet tothe outlet. 20) A heart pump according to any one of the claims 12 to19, wherein the controller is for: a) detecting movement of the impellercaused by a change in relative fluid pressures in the cavity portions;and, b) causing the magnetic bearing to control the axial position ofthe impeller to thereby alter the relative flow of fluid from the inletsto the outlets. 21) A heart pump according to any one of the claims 12to 20, wherein the controller is for: a) determining axial movement ofthe impeller away from a normal balance position; b) causing themagnetic bearing to move the impeller towards the normal position; c)monitoring the power used by the magnetic bearing; d) determining a newbalance position in accordance with the power used by the magneticbearing; and, e) causing the magnetic bearing to move the impeller tothe new balance position. 22) A heart pump according to claim 21,wherein the normal balance position is used to maintain required fluidflows from each inlet to each outlet. 23) A heart pump according toclaim 21 or claim 22, wherein the new balance position is offset fromthe normal balance position. 24) A heart pump according to claim 23,wherein the new balance position is used to adjust relative fluid flowsbetween the inlets and the outlets. 25) A heart pump according to anyone of the claims 12 to 24, wherein the indicator is determined using anindication of an electrical current used by the magnetic bearing. 26) Aheart pump according to claim 25, wherein the controller is fordetermining a rate of change of current used by the magnetic bearing tocause axial movement of the impeller. 27) A heart pump according to anyone of the claims 12 to 26, wherein the controller is for: a)determining movement of the impeller in a first axial direction; b)controlling the magnetic bearing to move the impeller in a second axialdirection opposite the first axial direction until at least one of: i)the power used by the magnetic bearing falls below a predeterminedamount; and, ii) the axial position of the impeller reaches a positionlimit. 28) A heart pump according to any one of the claims 12 to 27,wherein, the processing system includes: a) a memory for storinginstructions; and, b) a processor that executes the instructions,thereby causing the processor to: i) determine movement of the impellerin the first axial direction; ii) generate a signal for causing themagnetic bearing to move the impeller in the second axial direction;iii) determine an indicator indicative of the power used by the magneticbearing; and, iv) generate a signal for causing the magnetic bearing tocontrol the axial position of the impeller in accordance with theindicator to thereby control a fluid flow between the inlet and theoutlet. 29) A heart pump according to any one of the claims 1 to 28,wherein the heat pump includes a controller for: a) determining movementof an impeller from a balance position within a cavity, the cavityincluding at least one inlet and at least one outlet, and the impellerincluding vanes for urging fluid from the inlet to the outlet; b)causing a magnetic bearing to move the impeller to a new balanceposition based on an indication of power used by the magnetic bearing,the magnetic bearing including at least one coil for controlling anaxial position of the impeller within the cavity, and the new balanceposition being used to control fluid flow from the inlet to the outlet.30) A heart pump according to any one of the claims 1 to 29, wherein theheat pump includes a controller for controlling an axial position of animpeller within a cavity, a cavity including a first cavity portionhaving a first inlet and a first outlet and a second cavity portionhaving a second inlet and a second outlet, and the impeller includingfirst and second sets of vanes, each set of vanes being for urging fluidfrom a respective inlet to a respective outlet, the controllercontrolling the axial position such that if the relative pressure in thefirst cavity increases relative to the second cavity, the impeller ispositioned in the first cavity thereby increase the relative fluid flowsfrom the first outlet relative to the second outlet. 31) A heart pumpaccording to any one of the claims 1 to 30, wherein the heat pumpincludes a controller for controlling an axial position of an impellerwithin a cavity, a cavity including an inlet and an outlet, and theimpeller including vanes for urging fluid from the inlet to the outlet,the controller controlling the axial position such that if the pressurein the cavity increases, the impeller is moves away from the inlet,thereby reducing an outlet flow pressure. 32) A method of controlling aheart pump, the heart pump including: a) first and second cavityportions, each cavity portion including a respective inlet and outlet;b) a connecting tube extending between the first and second cavityportion; c) an impeller including: i) a first set of vanes mounted on afirst rotor in the first cavity portions; ii) a second set of vanesmounted on a second rotor in the second cavity portion; and, iii) ashaft connecting the first and second rotors, the shaft extendingthrough the connecting tube; d) a drive for rotating the impeller; and,e) a magnetic bearing including at least one bearing coil forcontrolling an axial position of the impeller, at least one of the driveand magnetic bearing being provided at least partially between the firstand second cavity portions, the method including: i) determiningmovement of the impeller in a first axial direction; ii) causing themagnetic bearing to move the impeller in a second axial directionopposite the first axial direction; iii) determining an indicatorindicative of the power used by the magnetic bearing; and, iv) causingthe magnetic bearing to control the axial position of the impeller inaccordance with the indicator to thereby control a fluid flow betweenthe inlets and the outlets. 33) A pump apparatus comprising: a) astator; b) an upper rotor arranged spaced apart above the stator; c) alower rotor arranged spaced apart under the stator; d) connecting meansprovided rotatably and vertically movable at the stator to connect theupper rotor and the lower rotor; e) first electromagnetic means that isprovided on one surface of the stator opposed to either one of permanentmagnets, which are provided on a bottom surface of the upper rotor and atop surface of the lower rotor, cooperates with the permanent magnet,and that generates an acting force with respect to the rotor in an axialdirection to thereby levitate the rotor; f) second electromagnetic meansthat is provided on the other surface of the stator opposed to the otherpermanent magnet, and that cooperates with the permanent magnet tothereby rotationally drive the rotor; g) first pumping means at which afirst impeller provided on a top surface of the upper rotor rotates in afirst pump chamber; and h) second pumping means at which a secondimpeller provided on a bottom surface of the lower rotor rotates in asecond pump chamber. 34) The pump apparatus according to claim 33,wherein the connecting means is composed of an axis of rotation insertedrotatably and vertically movable into a centre through-hole of thestator. 35) The pump apparatus according to claim 33 or 34, comprising acontrol section that moves the connecting means forward and backward inan axial direction and that can adjust a gap between the first impellerand an opposed surface of the first pump chamber, and a gap between thesecond impeller and the second pump chamber by coordinating them.